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Abstract:

A microfluidic device for detecting target nucleic acid sequences in a
sample, the microfluidic device having an array of hybridization
chambers, each of the hybridization chambers containing electrode pairs
for receiving an electrical pulse and electrochemiluminescent (ECL)
probes for hybridization with the target nucleic acid sequences, the ECL
probes being configured to emit a photon of light when hybridized with
one of the nucleic acid targets and excited by current between the
electrodes, wherein, the mass of the probes in each of the hybridization
chambers is less than 270 picograms.

Claims:

1. A microfluidic device for detecting target nucleic acid sequences in a
sample, the microfluidic device comprising: an array of hybridization
chambers, each of the hybridization chambers containing electrode pairs
for receiving an electrical pulse and electrochemiluminescent (ECL)
probes for hybridization with the target nucleic acid sequences, the ECL
probes being configured to emit a photon of light when hybridized with
one of the nucleic acid targets and excited by current between the
electrodes; wherein, the mass of the probes in each of the hybridization
chambers is less than 270 picograms.

2. The microfluidic test module according to claim 1 wherein the mass of
the probes in each of the hybridization chambers is less than 60
picograms.

3. The microfluidic test module according to claim 2 wherein the mass of
the probes in each of the hybridization chambers is less than 12
picograms.

4. The microfluidic test module according to claim 3 wherein the mass of
the probes in each of the hybridization chambers is less than 2.7
picograms.

5. A microfluidic device according to claim 1 wherein each of the
hybridization chambers contains one of the electrode pairs respectively.

6. A microfluidic device according to claim 5 wherein the hybridization
chambers each have a wall section that is optically transparent to the
light emitted by the ECL probes.

7. A microfluidic device according to claim 1 further comprising a
photosensor for detecting the light emitted by the ECL probes.

8. A microfluidic device according to claim 1 wherein the probes each
have an ECL luminophore that emits a photon when in an excited state, and
a functional moiety for quenching photon emission from the ECL
luminophore by resonant energy transfer.

9. A microfluidic device according to claim 8 wherein the probes are
configured such that the functional moiety for quenching photon emission
from the ECL luminophore is further from the ECL luminophore when the
probe forms a probe-target hybrid.

10. A microfluidic device according to claim 1 further comprising CMOS
circuitry configured to provide an electrical pulse to the electrodes.

11. A microfluidic device according to claim 10 wherein the electrical
pulse has a duration less than 0.69 seconds.

12. A microfluidic device according to claim 11 wherein the electrical
pulse has a current of 0.1 nanoamperes to 69.0 nanoamperes.

13. A microfluidic device according to claim 1 wherein the electrode
pairs have an anode and a cathode each having fingers configured such
that the fingers of the anode are interdigitated with the fingers of the
cathode.

14. A microfluidic device according to claim 13 wherein the anode and the
cathode are separated by a dielectric gap between 0.4 microns and 2.0
microns wide.

15. A microfluidic device according to claim 10 further comprising a cap
having reagent reservoirs for addition to the sample prior to detection
of the target sequences, wherein the electrodes and the probes are
between the cap and the CMOS circuitry.

16. A microfluidic device according to claim 15 wherein the reagent
reservoirs each have an outlet valve for retaining liquid reagent in the
reservoir until reagent addition to the sample is required.

17. A microfluidic device according to claim 7 wherein the photosensor is
an array of photodiodes positioned in registration with the hybridization
chambers such that each of the hybridization chambers corresponds to one
of the photodiodes respectively.

18. A microfluidic device according to claim 4 further comprising a
polymerase chain reaction (PCR) section for amplifying the target nucleic
acid sequences in the sample.

19. A microfluidic device according to claim 18 wherein the PCR section
has a heater element for thermal cycling the target nucleic acid
sequences with polymerase, the heater element being configured for
operative control by the CMOS circuitry.

20. A microfluidic device according to claim 19 further comprising a
plurality of sensors connected to the CMOS circuitry for feedback control
of the electrodes and the heater element.

Description:

FIELD OF THE INVENTION

[0001] The present invention relates to diagnostic devices that use
microsystems technologies (MST). In particular, the invention relates to
microfluidic and biochemical processing and analysis for molecular
diagnostics.

CO-PENDING APPLICATIONS

[0002] The following applications have been filed by the Applicant which
relate to the present application:

[0003] The disclosures of these co-pending applications are incorporated
herein by reference. The above applications have been identified by their
filing docket number, which will be substituted with the corresponding
application number, once assigned.

BACKGROUND OF THE INVENTION

[0004] Molecular diagnostics has emerged as a field that offers the
promise of early disease detection, potentially before symptoms have
manifested. Molecular diagnostic testing is used to detect: [0005]
Inherited disorders [0006] Acquired disorders [0007] Infectious diseases
[0008] Genetic predisposition to health-related conditions.

[0009] With high accuracy and fast turnaround times, molecular diagnostic
tests have the potential to reduce the occurrence of ineffective health
care services, enhance patient outcomes, improve disease management and
individualize patient care. Many of the techniques in molecular
diagnostics are based on the detection and identification of specific
nucleic acids, both deoxyribonucleic acid (DNA) and ribonucleic acid
(RNA), extracted and amplified from a biological specimen (such as blood
or saliva). The complementary nature of the nucleic acid bases allows
short sequences of synthesized DNA (oligonucleotides) to bond (hybridize)
to specific nucleic acid sequences for use in nucleic acid tests. If
hybridization occurs, then the complementary sequence is present in the
sample. This makes it possible, for example, to predict the disease a
person will contract in the future, determine the identity and virulence
of an infectious pathogen, or determine the response a person will have
to a drug.

Nucleic Acid Based Molecular Diagnostic Test

[0010] A nucleic acid based test has four distinct steps:

[0011] 1. Sample preparation

[0012] 2. Nucleic acid extraction

[0013] 3. Nucleic acid amplification (optional)

[0014] 4. Detection

[0015] Many sample types are used for genetic analysis, such as blood,
urine, sputum and tissue samples. The diagnostic test determines the type
of sample required as not all samples are representative of the disease
process. These samples have a variety of constituents, but usually only
one of these is of interest. For example, in blood, high concentrations
of erythrocytes can inhibit the detection of a pathogenic organism.
Therefore a purification and/or concentration step at the beginning of
the nucleic acid test is often required.

[0016] Blood is one of the more commonly sought sample types. It has three
major constituents: leukocytes (white blood cells), erythrocytes (red
blood cells) and thrombocytes (platelets). The thrombocytes facilitate
clotting and remain active in vitro. To inhibit coagulation, the specimen
is mixed with an agent such as ethylenediaminetetraacetic acid (EDTA)
prior to purification and concentration. Erythrocytes are usually removed
from the sample in order to concentrate the target cells. In humans,
erythrocytes account for approximately 99% of the cellular material but
do not carry DNA as they have no nucleus. Furthermore, erythrocytes
contain components such as haemoglobin that can interfere with the
downstream nucleic acid amplification process (described below). Removal
of erythrocytes can be achieved by differentially lysing the erythrocytes
in a lysis solution, leaving remaining cellular material intact which can
then be separated from the sample using centrifugation. This provides a
concentration of the target cells from which the nucleic acids are
extracted.

[0017] The exact protocol used to extract nucleic acids depends on the
sample and the diagnostic assay to be performed. For example, the
protocol for extracting viral RNA will vary considerably from the
protocol to extract genomic DNA. However, extracting nucleic acids from
target cells usually involves a cell lysis step followed by nucleic acid
purification. The cell lysis step disrupts the cell and nuclear
membranes, releasing the genetic material. This is often accomplished
using a lysis detergent, such as sodium dodecyl sulfate, which also
denatures the large amount of proteins present in the cells.

[0018] The nucleic acids are then purified with an alcohol precipitation
step, usually ice-cold ethanol or isopropanol, or via a solid phase
purification step, typically on a silica matrix in a column, resin or on
paramagnetic beads in the presence of high concentrations of a chaotropic
salt, prior to washing and then elution in a low ionic strength buffer.
An optional step prior to nucleic acid precipitation is the addition of a
protease which digests the proteins in order to further purify the
sample.

[0019] Other lysis methods include mechanical lysis via ultrasonic
vibration and thermal lysis where the sample is heated to 94° C.
to disrupt cell membranes.

[0020] The target DNA or RNA may be present in the extracted material in
very small amounts, particularly if the target is of pathogenic origin.
Nucleic acid amplification provides the ability to selectively amplify
(that is, replicate) specific targets present in low concentrations to
detectable levels.

[0021] The most commonly used nucleic acid amplification technique is the
polymerase chain reaction (PCR). PCR is well known in this field and
comprehensive description of this type of reaction is provided in E. van
Pelt-Verkuil et al., Principles and Technical Aspects of PCR
Amplification, Springer, 2008.

[0022] PCR is a powerful technique that amplifies a target DNA sequence
against a background of complex DNA. If RNA is to be amplified (by PCR),
it must be first transcribed into cDNA (complementary DNA) using an
enzyme called reverse transcriptase. Afterwards, the resulting cDNA is
amplified by PCR.

[0023] PCR is an exponential process that proceeds as long as the
conditions for sustaining the reaction are acceptable. The components of
the reaction are:

[0024] 1. pair of primers--short single strands of DNA with around 10-30
nucleotides complementary to the regions flanking the target sequence

[0025] 2. DNA polymerase--a thermostable enzyme that synthesizes DNA

[0026] 3. deoxyribonucleoside triphosphates (dNTPs)--provide the
nucleotides that are incorporated into the newly synthesized DNA strand

[0028] PCR typically involves placing these reactants in a small tube
(˜10-50 microlitres) containing the extracted nucleic acids. The
tube is placed in a thermal cycler; an instrument that subjects the
reaction to a series of different temperatures for varying amounts of
time. The standard protocol for each thermal cycle involves a
denaturation phase, an annealing phase, and an extension phase. The
extension phase is sometimes referred to as the primer extension phase.
In addition to such three-step protocols, two-step thermal protocols can
be employed, in which the annealing and extension phases are combined.
The denaturation phase typically involves raising the temperature of the
reaction to 90-95° C. to denature the DNA strands; in the
annealing phase, the temperature is lowered to ˜50-60° C.
for the primers to anneal; and then in the extension phase the
temperature is raised to the optimal DNA polymerase activity temperature
of 60-72° C. for primer extension. This process is repeated
cyclically around 20-40 times, the end result being the creation of
millions of copies of the target sequence between the primers.

[0029] There are a number of variants to the standard PCR protocol such as
multiplex PCR, linker-primed PCR, direct PCR, tandem PCR, real-time PCR
and reverse-transcriptase PCR, amongst others, which have been developed
for molecular diagnostics.

[0030] Multiplex PCR uses multiple primer sets within a single PCR mixture
to produce amplicons of varying sizes that are specific to different DNA
sequences. By targeting multiple genes at once, additional information
may be gained from a single test-run that otherwise would require several
experiments. Optimization of multiplex PCR is more difficult though and
requires selecting primers with similar annealing temperatures, and
amplicons with similar lengths and base composition to ensure the
amplification efficiency of each amplicon is equivalent.

[0031] Linker-primed PCR, also known as ligation adaptor PCR, is a method
used to enable nucleic acid amplification of essentially all DNA
sequences in a complex DNA mixture without the need for target-specific
primers. The method firstly involves digesting the target DNA population
with a suitable restriction endonuclease (enzyme). Double-stranded
oligonucleotide linkers (also called adaptors) with a suitable
overhanging end are then ligated to the ends of target DNA fragments
using a ligase enzyme. Nucleic acid amplification is subsequently
performed using oligonucleotide primers which are specific for the linker
sequences. In this way, all fragments of the DNA source which are flanked
by linker oligonucleotides can be amplified.

[0032] Direct PCR describes a system whereby PCR is performed directly on
a sample without any, or with minimal, nucleic acid extraction. It has
long been accepted that PCR reactions are inhibited by the presence of
many components of unpurified biological samples, such as the haem
component in blood. Traditionally, PCR has required extensive
purification of the target nucleic acid prior to preparation of the
reaction mixture. With appropriate changes to the chemistry and sample
concentration, however, it is possible to perform PCR with minimal DNA
purification, or direct PCR. Adjustments to the PCR chemistry for direct
PCR include increased buffer strength, the use of polymerases which have
high activity and processivity, and additives which chelate with
potential polymerase inhibitors.

[0033] Tandem PCR utilises two distinct rounds of nucleic acid
amplification to increase the probability that the correct amplicon is
amplified. One form of tandem PCR is nested PCR in which two pairs of PCR
primers are used to amplify a single locus in separate rounds of nucleic
acid amplification. The first pair of primers hybridize to the nucleic
acid sequence at regions external to the target nucleic acid sequence.
The second pair of primers (nested primers) used in the second round of
amplification bind within the first PCR product and produce a second PCR
product containing the target nucleic acid, that will be shorter than the
first one. The logic behind this strategy is that if the wrong locus were
amplified by mistake during the first round of nucleic acid
amplification, the probability is very low that it would also be
amplified a second time by a second pair of primers and thus ensures
specificity.

[0034] Real-time PCR, or quantitative PCR, is used to measure the quantity
of a PCR product in real time. By using a fluorophore-containing probe or
fluorescent dyes along with a set of standards in the reaction, it is
possible to quantitate the starting amount of nucleic acid in the sample.
This is particularly useful in molecular diagnostics where treatment
options may differ depending on the pathogen load in the sample.

[0035] Reverse-transcriptase PCR (RT-PCR) is used to amplify DNA from RNA.
Reverse transcriptase is an enzyme that reverse transcribes RNA into
complementary DNA (cDNA), which is then amplified by PCR. RT-PCR is
widely used in expression profiling, to determine the expression of a
gene or to identify the sequence of an RNA transcript, including
transcription start and termination sites. It is also used to amplify RNA
viruses such as human immunodeficiency virus or hepatitis C virus.

[0036] Isothermal amplification is another form of nucleic acid
amplification which does not rely on the thermal denaturation of the
target DNA during the amplification reaction and hence does not require
sophisticated machinery. Isothermal nucleic acid amplification methods
can therefore be carried out in primitive sites or operated easily
outside of a laboratory environment. A number of isothermal nucleic acid
amplification methods have been described, including Strand Displacement
Amplification, Transcription Mediated Amplification, Nucleic Acid
Sequence Based Amplification, Recombinase Polymerase Amplification,
Rolling Circle Amplification, Ramification Amplification,
Helicase-Dependent Isothermal DNA Amplification and Loop-Mediated
Isothermal Amplification.

[0037] Isothermal nucleic acid amplification methods do not rely on the
continuing heat denaturation of the template DNA to produce single
stranded molecules to serve as templates for further amplification, but
instead rely on alternative methods such as enzymatic nicking of DNA
molecules by specific restriction endonucleases, or the use of an enzyme
to separate the DNA strands, at a constant temperature.

[0038] Strand Displacement Amplification (SDA) relies on the ability of
certain restriction enzymes to nick the unmodified strand of
hemi-modified DNA and the ability of a 5'-3' exonuclease-deficient
polymerase to extend and displace the downstream strand. Exponential
nucleic acid amplification is then achieved by coupling sense and
antisense reactions in which strand displacement from the sense reaction
serves as a template for the antisense reaction. The use of nickase
enzymes which do not cut DNA in the traditional manner but produce a nick
on one of the DNA strands, such as N. Alw1, N. BstNB1 and Mly1, are
useful in this reaction. SDA has been improved by the use of a
combination of a heat-stable restriction enzyme (Ava1) and heat-stable
Exo-polymerase (Bst polymerase). This combination has been shown to
increase amplification efficiency of the reaction from 108 fold
amplification to 1010 fold amplification so that it is possible
using this technique to amplify unique single copy molecules.

[0039] Transcription Mediated Amplification (TMA) and Nucleic Acid
Sequence Based Amplification (NASBA) use an RNA polymerase to copy RNA
sequences but not corresponding genomic DNA. The technology uses two
primers and two or three enzymes, RNA polymerase, reverse transcriptase
and optionally RNase H (if the reverse transcriptase does not have RNase
activity). One primer contains a promoter sequence for RNA polymerase. In
the first step of nucleic acid amplification, this primer hybridizes to
the target ribosomal RNA (rRNA) at a defined site. Reverse transcriptase
creates a DNA copy of the target rRNA by extension from the 3' end of the
promoter primer. The RNA in the resulting RNA:DNA duplex is degraded by
the RNase activity of the reverse transcriptase if present or the
additional RNase H. Next, a second primer binds to the DNA copy. A new
strand of DNA is synthesized from the end of this primer by reverse
transcriptase, creating a double-stranded DNA molecule. RNA polymerase
recognizes the promoter sequence in the DNA template and initiates
transcription. Each of the newly synthesized RNA amplicons re-enters the
process and serves as a template for a new round of replication.

[0040] In Recombinase Polymerase Amplification (RPA), the isothermal
amplification of specific DNA fragments is achieved by the binding of
opposing oligonucleotide primers to template DNA and their extension by a
DNA polymerase. Heat is not required to denature the double-stranded DNA
(dsDNA) template. Instead, RPA employs recombinase-primer complexes to
scan dsDNA and facilitate strand exchange at cognate sites. The resulting
structures are stabilised by single-stranded DNA binding proteins
interacting with the displaced template strand, thus preventing the
ejection of the primer by branch migration. Recombinase disassembly
leaves the 3' end of the oligonucleotide accessible to a strand
displacing DNA polymerase, such as the large fragment of Bacillus
subtilis Pol I (Bsu), and primer extension ensues. Exponential nucleic
acid amplification is accomplished by the cyclic repetition of this
process.

[0041] Helicase-dependent amplification (HDA) mimics the in vivo system in
that it uses a DNA helicase enzyme to generate single-stranded templates
for primer hybridization and subsequent primer extension by a DNA
polymerase. In the first step of the HDA reaction, the helicase enzyme
traverses along the target DNA, disrupting the hydrogen bonds linking the
two strands which are then bound by single-stranded binding proteins.
Exposure of the single-stranded target region by the helicase allows
primers to anneal. The DNA polymerase then extends the 3' ends of each
primer using free deoxyribonucleoside triphosphates (dNTPs) to produce
two DNA replicates. The two replicated dsDNA strands independently enter
the next cycle of HDA, resulting in exponential nucleic acid
amplification of the target sequence.

[0042] Other DNA-based isothermal techniques include Rolling Circle
Amplification (RCA) in which a DNA polymerase extends a primer
continuously around a circular DNA template, generating a long DNA
product that consists of many repeated copies of the circle. By the end
of the reaction, the polymerase generates many thousands of copies of the
circular template, with the chain of copies tethered to the original
target DNA. This allows for spatial resolution of target and rapid
nucleic acid amplification of the signal. Up to 1012 copies of
template can be generated in 1 hour. Ramification amplification is a
variation of RCA and utilizes a closed circular probe (C-probe) or
padlock probe and a DNA polymerase with a high processivity to
exponentially amplify the C-probe under isothermal conditions.

[0043] Loop-mediated isothermal amplification (LAMP), offers high
selectivity and employs a DNA polymerase and a set of four specially
designed primers that recognize a total of six distinct sequences on the
target DNA. An inner primer containing sequences of the sense and
antisense strands of the target DNA initiates LAMP. The following strand
displacement DNA synthesis primed by an outer primer releases a
single-stranded DNA. This serves as template for DNA synthesis primed by
the second inner and outer primers that hybridize to the other end of the
target, which produces a stem-loop DNA structure. In subsequent LAMP
cycling one inner primer hybridizes to the loop on the product and
initiates displacement DNA synthesis, yielding the original stem-loop DNA
and a new stem-loop DNA with a stem twice as long. The cycling reaction
continues with accumulation of 109 copies of target in less than an
hour. The final products are stem-loop DNAs with several inverted repeats
of the target and cauliflower-like structures with multiple loops formed
by annealing between alternately inverted repeats of the target in the
same strand.

[0044] After completion of the nucleic acid amplification, the amplified
product must be analysed to determine whether the anticipated amplicon
(the amplified quantity of target nucleic acids) was generated. The
methods of analyzing the product range from simply determining the size
of the amplicon through gel electrophoresis, to identifying the
nucleotide composition of the amplicon using DNA hybridization.

[0045] Gel electrophoresis is one of the simplest ways to check whether
the nucleic acid amplification process generated the anticipated
amplicon. Gel electrophoresis uses an electric field applied to a gel
matrix to separate DNA fragments. The negatively charged DNA fragments
will move through the matrix at different rates, determined largely by
their size. After the electrophoresis is complete, the fragments in the
gel can be stained to make them visible. Ethidium bromide is a commonly
used stain which fluoresces under UV light.

[0046] The size of the fragments is determined by comparison with a DNA
size marker (a DNA ladder), which contains DNA fragments of known sizes,
run on the gel alongside the amplicon. Because the oligonucleotide
primers bind to specific sites flanking the target DNA, the size of the
amplified product can be anticipated and detected as a band of known size
on the gel. To be certain of the identity of the amplicon, or if several
amplicons have been generated, DNA probe hybridization to the amplicon is
commonly employed.

[0047] DNA hybridization refers to the formation of double-stranded DNA by
complementary base pairing. DNA hybridization for positive identification
of a specific amplification product requires the use of a DNA probe
around 20 nucleotides in length. If the probe has a sequence that is
complementary to the amplicon (target) DNA sequence, hybridization will
occur under favourable conditions of temperature, pH and ionic
concentration. If hybridization occurs, then the gene or DNA sequence of
interest was present in the original sample.

[0048] Optical detection is the most common method to detect
hybridization. Either the amplicons or the probes are labelled to emit
light through fluorescence or electrochemiluminescence. These processes
differ in the means of producing excited states of the light-producing
moieties, but both enable covalent labelling of nucleotide strands. In
electrochemiluminescence (ECL), light is produced by luminophore
molecules or complexes upon stimulation with an electric current. In
fluorescence, it is illumination with excitation light which leads to
emission.

[0049] Fluorescence is detected using an illumination source which
provides excitation light at a wavelength absorbed by the fluorescent
molecule, and a detection unit. The detection unit comprises a
photosensor (such as a photomultiplier tube or charge-coupled device
(CCD) array) to detect the emitted signal, and a mechanism (such as a
wavelength-selective filter) to prevent the excitation light from being
included in the photosensor output. The fluorescent molecules emit
Stokes-shifted light in response to the excitation light, and this
emitted light is collected by the detection unit. Stokes shift is the
frequency difference or wavelength difference between emitted light and
absorbed excitation light.

[0050] ECL emission is detected using a photosensor which is sensitive to
the emission wavelength of the ECL species being employed. For example,
transition metal-ligand complexes emit light at visible wavelengths, so
conventional photodiodes and CCDs are employed as photosensors. An
advantage of ECL is that, if ambient light is excluded, the ECL emission
can be the only light present in the detection system, which improves
sensitivity.

[0051] Microarrays allow for hundreds of thousands of DNA hybridization
experiments to be performed simultaneously. Microarrays are powerful
tools for molecular diagnostics with the potential to screen for
thousands of genetic diseases or detect the presence of numerous
infectious pathogens in a single test. A microarray consists of many
different DNA probes immobilized as spots on a substrate. The target DNA
(amplicon) is first labelled with a fluorescent or luminescent molecule
(either during or after nucleic acid amplification) and then applied to
the array of probes. The microarray is incubated in a temperature
controlled, humid environment for a number of hours or days while
hybridization between the probe and amplicon takes place. Following
incubation, the microarray must be washed in a series of buffers to
remove unbound strands. Once washed, the microarray surface is dried
using a stream of air (often nitrogen). The stringency of the
hybridization and washes is critical. Insufficient stringency can result
in a high degree of nonspecific binding. Excessive stringency can lead to
a failure of appropriate binding, which results in diminished
sensitivity. Hybridization is recognized by detecting light emission from
the labelled amplicons which have formed a hybrid with complementary
probes.

[0052] Fluorescence from microarrays is detected using a microarray
scanner which is generally a computer controlled inverted scanning
fluorescence confocal microscope which typically uses a laser for
excitation of the fluorescent dye and a photosensor (such as a
photomultiplier tube or CCD) to detect the emitted signal. The
fluorescent molecules emit Stokes-shifted light (described above) which
is collected by the detection unit.

[0053] The emitted fluorescence must be collected, separated from the
unabsorbed excitation wavelength, and transported to the detector. In
microarray scanners, a confocal arrangement is commonly used to eliminate
out-of-focus information by means of a confocal pinhole situated at an
image plane. This allows only the in-focus portion of the light to be
detected. Light from above and below the plane of focus of the object is
prevented from entering the detector, thereby increasing the signal to
noise ratio. The detected fluorescent photons are converted into
electrical energy by the detector which is subsequently converted to a
digital signal. This digital signal translates to a number representing
the intensity of fluorescence from a given pixel. Each feature of the
array is made up of one or more such pixels. The final result of a scan
is an image of the array surface. The exact sequence and position of
every probe on the microarray is known, and so the hybridized target
sequences can be identified and analysed simultaneously.

[0055] Despite the advantages that molecular diagnostic tests offer, the
growth of this type of testing in the clinical laboratory has been slower
than expected and remains a minor part of the practice of laboratory
medicine. This is primarily due to the complexity and costs associated
with nucleic acid testing compared with tests based on methods not
involving nucleic acids. The widespread adaptation of molecular
diagnostics testing to the clinical setting is intimately tied to the
development of instrumentation that significantly reduces the cost,
provides a rapid and automated assay from start (specimen processing) to
finish (generating a result) and operates without major intervention by
personnel.

[0056] A point-of-care technology serving the physician's office, the
hospital bedside or even consumer-based, at home, would offer many
advantages including: [0057] rapid availability of results enabling
immediate facilitation of treatment and improved quality of care. [0058]
ability to obtain laboratory values from testing very small samples.
[0059] reduced clinical workload. [0060] reduced laboratory workload and
improved office efficiency by reducing administrative work. [0061]
improved cost per patient through reduced length of stay of
hospitalization, conclusion of outpatient consultation at the first
visit, and reduced handling, storing and shipping of specimens. [0062]
facilitation of clinical management decisions such as infection control
and antibiotic use.

Lab-on-a-Chip (Loc) Based Molecular Diagnostics

[0063] Molecular diagnostic systems based on microfluidic technologies
provide the means to automate and speed up molecular diagnostic assays.
The quicker detection times are primarily due to the extremely low
volumes involved, automation, and the low-overhead inbuilt cascading of
the diagnostic process steps within a microfluidic device. Volumes in the
nanoliter and microliter scale also reduce reagent consumption and cost.
Lab-on-a-chip (LOC) devices are a common form of microfluidic device. LOC
devices have MST structures within a MST layer for fluid processing
integrated onto a single supporting substrate (usually silicon).
Fabrication using the VLSI (very large scale integrated) lithographic
techniques of the semiconductor industry keeps the unit cost of each LOC
device very low. However, controlling fluid flow through the LOC device,
adding reagents, controlling reaction conditions and so on necessitate
bulky external plumbing and electronics. Connecting a LOC device to these
external devices effectively restricts the use of LOC devices for
molecular diagnostics to the laboratory setting. The cost of the external
equipment and complexity of its operation precludes LOC-based molecular
diagnostics as a practical option for point-of-care settings.

[0064] In view of the above, there is a need for a molecular diagnostic
system based on a LOC device for use at point-of-care.

[0066] an array of hybridization chambers, each of the hybridization
chambers containing electrode pairs for receiving an electrical pulse and
electrochemiluminescent (ECL) probes for hybridization with the target
nucleic acid sequences, the ECL probes being configured to emit a photon
of light when hybridized with one of the nucleic acid targets and excited
by current between the electrodes; wherein,

[0067] the mass of the probes in each of the hybridization chambers is
less than 270 picograms.

[0068] Preferably, the mass of the probes in each of the hybridization
chambers is less than 60 picograms.

[0069] Preferably, the mass of the probes in each of the hybridization
chambers is less than 12 picograms.

[0070] Preferably, the mass of the probes in each of the hybridization
chambers is less than 2.7 picograms.

[0071] Preferably, each of the hybridization chambers contains one of the
electrode pairs respectively.

[0072] Preferably, the hybridization chambers each have a wall section
that is optically transparent to the light emitted by the ECL probes.

[0073] Preferably, the microfluidic device also has a photosensor for
detecting the light emitted by the ECL probes.

[0074] Preferably, the probes each have an ECL luminophore that emits a
photon when in an excited state, and a functional moiety for quenching
photon emission from the ECL luminophore by resonant energy transfer.

[0075] Preferably, the probes are configured such that the functional
moiety for quenching photon emission from the ECL luminophore is further
from the ECL luminophore when the probe forms a probe-target hybrid.

[0076] Preferably, the microfluidic device also has CMOS circuitry
configured to provide an electrical pulse to the electrodes.

[0077] Preferably, the electrical pulse has a duration less than 0.69
seconds.

[0078] Preferably, the electrical pulse has a current of 0.1 nanoamperes
to 69.0 nanoamperes.

[0079] Preferably, the electrode pairs have an anode and a cathode each
having fingers configured such that the fingers of the anode are
interdigitated with the fingers of the cathode.

[0080] Preferably, the anode and the cathode are separated by a dielectric
gap between 0.4 microns and 2.0 microns wide.

[0081] Preferably, the microfluidic device also has a cap having reagent
reservoirs for addition to the sample prior to detection of the target
sequences, wherein the electrodes and the probes are between the cap and
the CMOS circuitry.

[0082] Preferably, the reagent reservoirs each have an outlet valve for
retaining liquid reagent in the reservoir until reagent addition to the
sample is required.

[0083] Preferably, the photosensor is an array of photodiodes positioned
in registration with the hybridization chambers such that each of the
hybridization chambers corresponds to one of the photodiodes
respectively.

[0084] Preferably, the microfluidic device also has a polymerase chain
reaction (PCR) section for amplifying the target nucleic acid sequences
in the sample.

[0085] Preferably, the PCR section has a heater element for thermal
cycling the target nucleic acid sequences with polymerase, the heater
element being configured for operative control by the CMOS circuitry.

[0086] Preferably, the microfluidic device also has a plurality of sensors
connected to the CMOS circuitry for feedback control of the electrodes
and the heater element.

[0087] The low probe volume provides for low probe cost, in turn,
permitting the inexpensive assay system. The
electrochemiluminescence-based assay target detection obviates any need,
of the assay system, for an excitation light source, excitation optics,
and optical filter elements, in turn, providing for a more compact and
more inexpensive assay system. The absence of the requirement for the
rejection of any excitation light also simplifies the detector circuitry,
making the assay system even more inexpensive.

BRIEF DESCRIPTION OF THE DRAWINGS

[0088] Preferred embodiments of the present invention will now be
described by way of example only with reference to the accompanying
drawings, in which:

[0225] FIG. 137 schematically illustrates the positive control luminescent
probe of FIG. 136 in its open configuration;

[0226] FIG. 138 is an enlarged view of the hybridization chamber of LOC
variant L;

[0227] FIG. 139 is an enlarged view of the hybridization chamber array of
LOC variant L showing the distribution of calibration chambers;

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

Overview

[0228] This overview identifies the main components of a molecular
diagnostic system that incorporates embodiments of the present invention.
Comprehensive details of the system architecture and operation are set
out later in the specification.

[0229] Referring to FIGS. 1, 2, 3, 104 and 105, the system has the
following top level components:

[0230] Test modules 10 and 11 are the size of a typical USB memory key and
very cheap to produce. Test modules 10 and 11 each contain a microfluidic
device, typically in the form of a lab-on-a-chip (LOC) device 30
preloaded with reagents and typically more than 1000 probes for the
molecular diagnostic assay (see FIGS. 1 and 104). Test module 10
schematically shown in FIG. 1 uses a fluorescence-based detection
technique to identify target molecules, while test module 11 in FIG. 104
uses an electrochemiluminescence-based detection technique. The LOC
device 30 has an integrated photosensor 44 for fluorescence or
electrochemiluminescence detection (described in detail below). Both test
modules 10 and 11 use a standard Micro-USB plug 14 for power, data and
control, both have a printed circuit board (PCB) 57, and both have
external power supply capacitors 32 and an inductor 15. The test modules
10 and 11 are both single-use only for mass production and distribution
in sterile packaging ready for use.

[0231] The outer casing 13 has a macroreceptacle 24 for receiving the
biological sample and a removable sterile sealing tape 22, preferably
with a low tack adhesive, to cover the macroreceptacle prior to use. A
membrane seal 408 with a membrane guard 410 forms part of the outer
casing 13 to reduce dehumidification within the test module while
providing pressure relief from small air pressure fluctuations. The
membrane guard 410 protects the membrane seal 408 from damage.

[0232] Test module reader 12 powers the test module 10 or 11 via Micro-USB
port 16. The test module reader 12 can adopt many different forms and a
selection of these are described later. The version of the reader 12
shown in FIGS. 1, 3 and 104 is a smart phone embodiment. A block diagram
of this reader 12 is shown in FIG. 3. Processor 42 runs application
software from program storage 43. The processor 42 also interfaces with
the display screen 18 and user interface (UI) touch screen 17 and buttons
19, a cellular radio 21, wireless network connection 23, and a satellite
navigation system 25. The cellular radio 21 and wireless network
connection 23 are used for communications. Satellite navigation system 25
is used for updating epidemiological databases with location data. The
location data can, alternatively, be entered manually via the touch
screen 17 or buttons 19. Data storage 27 holds genetic and diagnostic
information, test results, patient information, assay and probe data for
identifying each probe and its array position. Data storage 27 and
program storage 43 may be shared in a common memory facility. Application
software installed on the test module reader 12 provides analysis of
results, along with additional test and diagnostic information.

[0233] To conduct a diagnostic test, the test module 10 (or test module
11) is inserted into the Micro-USB port 16 on the test module reader 12.
The sterile sealing tape 22 is peeled back and the biological sample (in
a liquid form) is loaded into the sample macroreceptacle 24. Pressing
start button 20 initiates testing via the application software. The
sample flows into the LOC device 30 and the on-board assay extracts,
incubates, amplifies and hybridizes the sample nucleic acids (the target)
with presynthesized hybridization-responsive oligonucleotide probes. In
the case of test module 10 (which uses fluorescence-based detection), the
probes are fluorescently labelled and the LED 26 housed in the casing 13
provides the necessary excitation light to induce fluorescence emission
from the hybridized probes (see FIGS. 1 and 2). In test module 11 (which
uses electrochemiluminescence (ECL) detection), the LOC device 30 is
loaded with ECL probes (discussed above) and the LED 26 is not necessary
for generating the luminescent emission. Instead, electrodes 860 and 870
provide the excitation electrical current (see FIG. 105). The emission
(fluorescent or luminescent) is detected using a photosensor 44
integrated into CMOS circuitry of each LOC device. The detected signal is
amplified and converted to a digital output which is analyzed by the test
module reader 12. The reader then displays the results.

[0234] The data may be saved locally and/or uploaded to a network server
containing patient records. The test module 10 or 11 is removed from the
test module reader 12 and disposed of appropriately.

[0235] FIGS. 1, 3 and 104 show the test module reader 12 configured as a
mobile phone/smart phone 28. In other forms, the test module reader is a
laptop/notebook 101, a dedicated reader 103, an ebook reader 107, a
tablet computer 109 or desktop computer 105 for use in hospitals, private
practices or laboratories (see FIG. 106). The reader can interface with a
range of additional applications such as patient records, billing, online
databases and multi-user environments. It can also be interfaced with a
range of local or remote peripherals such as printers and patient smart
cards.

[0236] Referring to FIG. 107, the data generated by the test module 10 can
be used to update, via the reader 12 and network 125, the epidemiological
databases hosted on the hosting system for epidemiological data 111, the
genetic databases hosted on the hosting system for genetic data 113, the
electronic health records hosted on the hosting system for electronic
health records (EHR) 115, the electronic medical records hosted on the
hosting system for electronic medical records (EMR) 121, and the personal
health records hosted on the hosting system for personal health records
(PHR) 123. Conversely, the epidemiological data hosted on the hosting
system for epidemiological data 111, the genetic data hosted on the
hosting system for genetic data 113, the electronic health records hosted
on the hosting system for electronic health records (EHR) 115, the
electronic medical records hosted on the hosting system for electronic
medical records (EMR) 121, and the personal health records hosted on the
hosting system for personal health records (PHR) 123, can be used to
update, via network 125 and the reader 12, the digital memory in the LOC
30 of the test module 10.

[0237] Referring back to FIGS. 1, 2, 104 and 105 the reader 12 uses
battery power in the mobile phone configuration. The mobile phone reader
contains all test and diagnostic information preloaded. Data can also be
loaded or updated via a number of wireless or contact interfaces to
enable communications with peripheral devices, computers or online
servers. A Micro-USB port 16 is provided for connection to a computer or
mains power supply for battery recharge.

[0238] FIG. 63 shows an embodiment of the test module 10 used for tests
that only require a positive or negative result for a particular target,
such as testing whether a person is infected with, for example, H1N1
Influenza A virus. Only a purpose built USB power/indicator-only module
47 is adequate. No other reader or application software is necessary. An
indicator 45 on the USB power/indicator-only module 47 signals positive
or negative results. This configuration is well suited to mass screening.

[0239] Additional items supplied with the system may include a test tube
containing reagents for pre-treatment of certain samples, along with
spatula and lancet for sample collection. FIG. 63 shows an embodiment of
the test module incorporating a spring-loaded, retractable lancet 390 and
lancet release button 392 for convenience. A satellite phone can be used
in remote areas.

Test Module Electronics

[0240] FIGS. 2 and 105 are block diagrams of the electronic components in
the test modules 10 and 11, respectively. The CMOS circuitry integrated
in the LOC device 30 has a USB device driver 36, a controller 34, a
USB-compatible LED driver 29, clock 33, power conditioner 31, RAM 38 and
program and data flash memory 40. These provide the control and memory
for the entire test module 10 or 11 including the photosensor 44, the
temperature sensors 170, the liquid sensors 174, and the various heaters
152, 154, 182, 234, together with associated drivers 37 and 39 and
registers 35 and 41. Only the LED 26 (in the case of test module 10),
external power supply capacitors 32 and the Micro-USB plug 14 are
external to the LOC device 30. The LOC devices 30 include bond-pads for
making connections to these external components. The RAM 38 and the
program and data flash memory 40 have the application software and the
diagnostic and test information (Flash/Secure storage, e.g. via
encryption) for over 1000 probes. In the case of test module 11
configured for ECL detection, there is no LED 26 (see FIGS. 104 and 105).
Data is encrypted by the LOC device 30 for secure storage and secure
communication with an external device. The LOC devices 30 are loaded with
electrochemiluminescent probes and the hybridization chambers each have a
pair of ECL excitation electrodes 860 and 870.

[0241] Many types of test modules 10 are manufactured in a number of test
forms, ready for off-the-shelf use. The differences between the test
forms lie in the on board assay of reagents and probes.

[0278] The above lists are not exhaustive and the diagnostic system can be
configured to detect a much greater variety of diseases and conditions
using nucleic acid and proteomic analysis.

DETAILED ARCHITECTURE OF SYSTEM COMPONENTS

LOC Device

[0279] The LOC device 30 is central to the diagnostic system. It rapidly
performs the four major steps of a nucleic acid based molecular
diagnostic assay, i.e. sample preparation, nucleic acid extraction,
nucleic acid amplification, and detection, using a microfluidic platform.
The LOC device also has alternative uses, and these are detailed later.
As discussed above, test modules 10 and 11 can adopt many different
configurations to detect different targets Likewise, the LOC device 30
has numerous different embodiments tailored to the target(s) of interest.
One form of the LOC device 30 is LOC device 301 for fluorescent detection
of target nucleic acid sequences in the pathogens of a whole blood
sample. For the purposes of illustration, the structure and operation of
LOC device 301 is now described in detail with reference to FIGS. 4 to 26
and 27 to 57.

[0280] FIG. 4 is a schematic representation of the architecture of the LOC
device 301. For convenience, process stages shown in FIG. 4 are indicated
with the reference numeral corresponding to the functional sections of
the LOC device 301 that perform that process stage. The process stages
associated with each of the major steps of a nucleic acid based molecular
diagnostic assay are also indicated: sample input and preparation 288,
extraction 290, incubation 291, amplification 292 and detection 294. The
various reservoirs, chambers, valves and other components of the LOC
device 301 will be described in more detail later.

[0281] FIG. 5 is a perspective view of the LOC device 301. It is
fabricated using high volume CMOS and MST (microsystems technology)
manufacturing techniques. The laminar structure of the LOC device 301 is
illustrated in the schematic (not to scale) partial section view of FIG.
12. The LOC device 301 has a silicon substrate 84 which supports the
CMOS+MST chip 48, comprising CMOS circuitry 86 and an MST layer 87, with
a cap 46 overlaying the MST layer 87. For the purposes of this patent
specification, the term `MST layer` is a reference to a collection of
structures and layers that process the sample with various reagents.
Accordingly, these structures and components are configured to define
flow-paths with characteristic dimensions that will support capillary
driven flow of liquids with physical characteristics similar to those of
the sample during processing. In light of this, the MST layer and
components are typically fabricated using surface micromachining
techniques and/or bulk micromachining techniques. However, other
fabrication methods can also produce structures and components
dimensioned for capillary driven flows and processing very small volumes.
The specific embodiments described in this specification show the MST
layer as the structures and active components supported on the CMOS
circuitry 86, but excluding the features of the cap 46. However, the
skilled addressee will appreciate that the MST layer need not have
underlying CMOS or indeed an overlying cap in order for it to process the
sample.

[0282] The overall dimensions of the LOC device shown in the following
figures are 1760 μm×5824 μm. Of course, LOC devices
fabricated for different applications may have different dimensions.

[0283] FIG. 6 shows the features of the MST layer 87 superimposed with the
features of the cap. Insets AA to AD, AG and AH shown in FIG. 6 are
enlarged in FIGS. 13, 14, 35, 56, 55 and 58, respectively, and described
in detail below for a comprehensive understanding of each structure
within the LOC device 301. FIGS. 7 to 10 show the features of the cap 46
in isolation while FIG. 11 shows the CMOS+MST device 48 structures in
isolation.

Laminar Structure

[0284] FIGS. 12 and 22 are sketches that diagrammatically show the laminar
structure of the CMOS+MST device 48, the cap 46 and the fluidic
interaction between the two. The figures are not to scale for the
purposes of illustration. FIG. 12 is a schematic section view through the
sample inlet 68 and FIG. 22 is a schematic section through the reservoir
54. As best shown in FIG. 12, the CMOS+MST device 48 has a silicon
substrate 84 which supports the CMOS circuitry 86 that operates the
active elements within the MST layer 87 above. A passivation layer 88
seals and protects the CMOS layer 86 from the fluid flows through the MST
layer 87.

[0285] Fluid flows through both the cap channels 94 and the MST channels
90 (see for example FIGS. 7 and 16) in the cap layer 46 and MST channel
layer 100, respectively. Cell transport occurs in the larger channels 94
fabricated in the cap 46, while biochemical processes are carried out in
the smaller MST channels 90. Cell transport channels are sized so as to
be able to transport cells in the sample to predetermined sites in the
MST channels 90. Transportation of cells with sizes greater than 20
microns (for example, certain leukocytes) requires channel dimensions
greater than 20 microns, and therefore a cross sectional area transverse
to the flow of greater than 400 square microns. MST channels,
particularly at locations in the LOC where transport of cells is not
required, can be significantly smaller.

[0286] It will be appreciated that cap channel 94 and MST channel 90 are
generic references and particular MST channels 90 may also be referred to
as (for example) heated microchannels or dialysis MST channels in light
of their particular function. MST channels 90 are formed by etching
through a MST channel layer 100 deposited on the passivation layer 88 and
patterned with photoresist. The MST channels 90 are enclosed by a roof
layer 66 which forms the top (with respect to the orientation shown in
the figures) of the CMOS+MST device 48.

[0287] Despite sometimes being shown as separate layers, the cap channel
layer 80 and the reservoir layer 78 are formed from a unitary piece of
material. Of course, the piece of material may also be non-unitary. This
piece of material is etched from both sides in order to form a cap
channel layer 80 in which the cap channels 94 are etched and the
reservoir layer 78 in which the reservoirs 54, 56, 58, 60 and 62 are
etched. Alternatively, the reservoirs and the cap channels are formed by
a micromolding process. Both etching and micromolding techniques are used
to produce channels with cross sectional areas transverse to the flow as
large as 20,000 square microns, and as small as 8 square microns.

[0288] At different locations in the LOC device, there can be a range of
appropriate choices for the cross sectional area of the channel
transverse to the flow. Where large quantities of sample, or samples with
large constituents, are contained in the channel, a cross-sectional area
of up to 20,000 square microns (for example, a 200 micron wide channel in
a 100 micron thick layer) is suitable. Where small quantities of liquid,
or mixtures without large cells present, are contained in the channel, a
very small cross sectional area transverse to the flow is preferable.

[0296] The cap 46 and the CMOS+MST layers 48 are in fluid communication
via corresponding openings in the lower seal 64 and the roof layer 66.
These openings are referred to as uptakes 96 and downtakes 92 depending
on whether fluid is flowing from the MST channels 90 to the cap channels
94 or vice versa.

LOC Device Operation

[0297] The operation of the LOC device 301 is described below in a
step-wise fashion with reference to analysing pathogenic DNA in a blood
sample. Of course, other types of biological or non-biological fluid are
also analysed using an appropriate set, or combination, of reagents, test
protocols, LOC variants and detection systems. Referring back to FIG. 4,
there are five major steps involved in analysing a biological sample,
comprising sample input and preparation 288, nucleic acid extraction 290,
nucleic acid incubation 291, nucleic acid amplification 292 and detection
and analysis 294.

[0298] The sample input and preparation step 288 involves mixing the blood
with an anticoagulant 116 and then separating pathogens from the
leukocytes and erythrocytes with the pathogen dialysis section 70. As
best shown in FIGS. 7 and 12, the blood sample enters the device via the
sample inlet 68. Capillary action draws the blood sample along the cap
channel 94 to the reservoir 54. Anticoagulant is released from the
reservoir 54 as the sample blood flow opens its surface tension valve 118
(see FIGS. 15 and 22). The anticoagulant prevents the formation of clots
which would block the flow.

[0299] As best shown in FIG. 22, the anticoagulant 116 is drawn out of the
reservoir 54 by capillary action and into the MST channel 90 via the
downtake 92. The downtake 92 has a capillary initiation feature (CIF) 102
to shape the geometry of the meniscus such that it does not anchor to the
rim of the downtake 92. Vent holes 122 in the upper seal 82 allows air to
replace the anticoagulant 116 as it is drawn out of the reservoir 54.

[0300] The MST channel 90 shown in FIG. 22 is part of a surface tension
valve 118. The anticoagulant 116 fills the surface tension valve 118 and
pins a meniscus 120 to the uptake 96 to a meniscus anchor 98. Prior to
use, the meniscus 120 remains pinned at the uptake 96 so the
anticoagulant does not flow into the cap channel 94. When the blood flows
through the cap channel 94 to the uptake 96, the meniscus 120 is removed
and the anticoagulant is drawn into the flow.

[0301] FIGS. 15 to 21 show Inset AE which is a portion of Inset AA shown
in FIG. 13. As shown in FIGS. 15, 16 and 17, the surface tension valve
118 has three separate MST channels 90 extending between respective
downtakes 92 and uptakes 96. The number of MST channels 90 in a surface
tension valve can be varied to change the flow rate of the reagent into
the sample mixture. As the sample mixture and the reagents mix together
by diffusion, the flow rate out of the reservoir determines the
concentration of the reagent in the sample flow. Hence, the surface
tension valve for each of the reservoirs is configured to match the
desired reagent concentration.

[0302] The blood passes into a pathogen dialysis section 70 (see FIGS. 4
and 15) where target cells are concentrated from the sample using an
array of apertures 164 sized according to a predetermined threshold.
Cells smaller than the threshold pass through the apertures while larger
cells do not pass through the apertures. Unwanted cells, which may be
either the larger cells withheld by the array of apertures 164 or the
smaller cells that pass through the apertures, are redirected to a waste
unit 76 while the target cells continue as part of the assay.

[0303] In the pathogen dialysis section 70 described here, the pathogens
from the whole blood sample are concentrated for microbial DNA analysis.
The array of apertures is formed by a multitude of 3 micron diameter
holes 164 fluidically connecting the input flow in the cap channel 94 to
a target channel 74. The 3 micron diameter apertures 164 and the dialysis
uptake holes 168 for the target channel 74 are connected by a series of
dialysis MST channels 204 (best shown in FIGS. 15 and 21). Pathogens are
small enough to pass through the 3 micron diameter apertures 164 and fill
the target channel 74 via the dialysis MST channels 204. Cells larger
than 3 microns, such as erythrocytes and leukocytes, stay in the waste
channel 72 in the cap 46 which leads to a waste reservoir 76 (see FIG.
7).

[0304] Other aperture shapes, sizes and aspect ratios can be used to
isolate specific pathogens or other target cells such as leukocytes for
human DNA analysis. Greater detail on the dialysis section and dialysis
variants is provided later.

[0305] Referring again to FIGS. 6 and 7, the flow is drawn through the
target channel 74 to the surface tension valve 128 of the lysis reagent
reservoir 56. The surface tension valve 128 has seven MST channels 90
extending between the lysis reagent reservoir 56 and the target channel
74. When the menisci are unpinned by the sample flow, the flow rate from
all seven of the MST channels 90 will be greater than the flow rate from
the anticoagulant reservoir 54 where the surface tension valve 118 has
three MST channels 90 (assuming the physical characteristics of the
fluids are roughly equivalent). Hence the proportion of lysis reagent in
the sample mixture is greater than that of the anticoagulant.

[0306] The lysis reagent and target cells mix by diffusion in the target
channel 74 within the chemical lysis section 130. A boiling-initiated
valve 126 stops the flow until sufficient time has passed for diffusion
and lysis to take place, releasing the genetic material from the target
cells (see FIGS. 6 and 7). The structure and operation of the
boiling-initiated valves are described in greater detail below with
reference to FIGS. 31 and 32. Other active valve types (as opposed to
passive valves such as the surface tension valve 118) have also been
developed by the Applicant which may be used here instead of the
boiling-initiated valve. These alternative valve designs are also
described later.

[0307] When the boiling-initiated valve 126 opens, the lysed cells flow
into a mixing section 131 for pre-amplification restriction digestion and
linker ligation.

[0308] Referring to FIG. 13, restriction enzymes, linkers and ligase are
released from the reservoir 58 when the flow unpins the menisci at the
surface tension valve 132 at the start of the mixing section 131. The
mixture flows the length of the mixing section 131 for diffusion mixing.
At the end of the mixing section 131 is downtake 134 leading into the
incubator inlet channel 133 of the incubation section 114 (see FIG. 13).
The incubator inlet channel 133 feeds the mixture into a serpentine
configuration of heated microchannels 210 which provides an incubation
chamber for holding the sample during restriction digestion and ligation
of the linkers (see FIGS. 13 and 14).

[0309] FIGS. 23, 24, 25, 26, 27, 28 and 29 show the layers of the LOC
device 301 within Inset AB of FIG. 6. Each figure shows the sequential
addition of layers forming the structures of the CMOS+MST layer 48 and
the cap 46. Inset AB shows the end of the incubation section 114 and the
start of the amplification section 112. As best shown in FIGS. 14 and 23,
the flow fills the microchannels 210 of the incubation section 114 until
reaching the boiling-initiated valve 106 where the flow stops while
diffusion takes place. As discussed above, the microchannel 210 upstream
of the boiling-initiated valve 106 becomes an incubation chamber
containing the sample, restriction enzymes, ligase and linkers. The
heaters 154 are then activated and held at constant temperature for a
specified time for restriction digestion and linker ligation to occur.

[0310] The skilled worker will appreciate that this incubation step 291
(see FIG. 4) is optional and only required for some nucleic acid
amplification assay types. Furthermore, in some instances, it may be
necessary to have a heating step at the end of the incubation period to
spike the temperature above the incubation temperature. The temperature
spike inactivates the restriction enzymes and ligase prior to entering
the amplification section 112. Inactivation of the restriction enzymes
and ligase has particular relevance when isothermal nucleic acid
amplification is being employed.

[0311] Following incubation, the boiling-initiated valve 106 is activated
(opened) and the flow resumes into the amplification section 112.
Referring to FIGS. 31 and 32, the mixture fills the serpentine
configuration of heated microchannels 158, which form one or more
amplification chambers, until it reaches the boiling-initiated valve 108.
As best shown in the schematic section view of FIG. 30, amplification mix
(dNTPs, primers, buffer) is released from reservoir 60 and polymerase is
subsequently released from reservoir 62 into the intermediate MST channel
212 connecting the incubation and amplification sections (114 and 112
respectively).

[0312] FIGS. 35 to 51 show the layers of the LOC device 301 within Inset
AC of FIG. 6. Each figure shows the sequential addition of layers forming
the structures of the CMOS+MST device 48 and the cap 46. Inset AC is at
the end of the amplification section 112 and the start of the
hybridization and detection section 52. The incubated sample,
amplification mix and polymerase flow through the microchannels 158 to
the boiling-initiated valve 108. After sufficient time for diffusion
mixing, the heaters 154 in the microchannels 158 are activated for
thermal cycling or isothermal amplification. The amplification mix goes
through a predetermined number of thermal cycles or a preset
amplification time to amplify sufficient target DNA. After the nucleic
acid amplification process, the boiling-initiated valve 108 opens and
flow resumes into the hybridization and detection section 52. The
operation of boiling-initiated valves is described in more detail later.

[0313] As shown in FIG. 52, the hybridization and detection section 52 has
an array of hybridization chambers 110. FIGS. 52, 53, 54 and 56 show the
hybridization chamber array 110 and individual hybridization chambers 180
in detail. At the entrance to the hybridization chamber 180 is a
diffusion barrier 175 which prevents diffusion of the target nucleic
acid, probe strands and hybridized probes between the hybridization
chambers 180 during hybridization so as to prevent erroneous
hybridization detection results. The diffusion barriers 175 present a
flow-path-length that is long enough to prevent the target sequences and
probes diffusing out of one chamber and contaminating another chamber
within the time taken for the probes and nucleic acids to hybridize and
the signal to be detected, thus avoiding an erroneous result.

[0314] Another mechanism to prevent erroneous readings is to have
identical probes in a number of the hybridization chambers. The CMOS
circuitry 86 derives a single result from the photodiodes 184
corresponding to the hybridization chambers 180 that contain identical
probes. Anomalous results can be disregarded or weighted differently in
the derivation of the single result.

[0315] The thermal energy required for hybridization is provided by
CMOS-controlled heaters 182 (described in more detail below). After the
heater is activated, hybridization occurs between complementary
target-probe sequences. The LED driver 29 in the CMOS circuitry 86
signals the LED 26 located in the test module 10 to illuminate. These
probes only fluoresce when hybridization has occurred thereby avoiding
washing and drying steps that are typically required to remove unbound
strands. Hybridization forces the stem-and-loop structure of the FRET
probes 186 to open, which allows the fluorophore to emit fluorescent
energy in response to the LED excitation light, as discussed in greater
detail later. Fluorescence is detected by a photodiode 184 in the CMOS
circuitry 86 underlying each hybridization chamber 180 (see hybridization
chamber description below). The photodiodes 184 for all hybridization
chambers and associated electronics collectively form the photosensor 44
(see FIG. 60). In other embodiments, the photosensor may be an array of
charge coupled devices (CCD array). The detected signal from the
photodiodes 184 is amplified and converted to a digital output which is
analyzed by the test module reader 12. Further details of the detection
method are described later.

Additional Details for the LOC Device

Modularity of the Design

[0316] The LOC device 301 has many functional sections, including the
reagent reservoirs 54, 56, 58, 60 and 62, the dialysis section 70, lysis
section 130, incubation section 114, and amplification section 112, valve
types, the humidifier and humidity sensor. In other embodiments of the
LOC device, these functional sections can be omitted, additional
functional sections can be added or the functional sections can be used
for alternative purposes to those described above.

[0317] For example, the incubation section 114 can be used as the first
amplification section 112 of a tandem amplification assay system, with
the chemical lysis reagent reservoir 56 being used to add the first
amplification mix of primers, dNTPs and buffer and reagent reservoir 58
being used for adding the reverse transcriptase and/or polymerase. A
chemical lysis reagent can also be added to the reservoir 56 along with
the amplification mix if chemical lysis of the sample is desired or,
alternatively, thermal lysis can occur in the incubation section by
heating the sample for a predetermined time. In some embodiments, an
additional reservoir can be incorporated immediately upstream of
reservoir 58 for the mix of primers, dNTPs and buffer if there is a
requirement for chemical lysis and a separation of this mix from the
chemical lysis reagent is desired.

[0318] In some circumstances it may be desirable to omit a step, such as
the incubation step 291. In this case, a LOC device can be specifically
fabricated to omit the reagent reservoir 58 and incubation section 114,
or the reservoir can simply not be loaded with reagents or the active
valves, if present, not activated to dispense the reagents into the
sample flow, and the incubation section then simply becomes a channel to
transport the sample from the lysis section 130 to the amplification
section 112. The heaters are independently operable and therefore, where
reactions are dependent on heat, such as thermal lysis, programming the
heaters not to activate during this step ensures thermal lysis does not
occur in LOC devices that do not require it. The dialysis section 70 can
be located at the beginning of the fluidic system within the microfluidic
device as shown in FIG. 4 or can be located anywhere else within the
microfluidic device. For example, dialysis after the amplification phase
292 to remove cellular debris prior to the hybridization and detection
step 294 may be beneficial in some circumstances. Alternatively, two or
more dialysis sections can be incorporated at any location throughout the
LOC device. Similarly, it is possible to incorporate additional
amplification sections 112 to enable multiple targets to be amplified in
parallel or in series prior to being detected in the hybridization
chamber arrays 110 with specific nucleic acid probes. For analysis of
samples like whole blood, in which dialysis is not required, the dialysis
section 70 is simply omitted from the sample input and preparation
section 288 of the LOC design. In some cases, it is not necessary to omit
the dialysis section 70 from the LOC device even if the analysis does not
require dialysis. If there is no geometric hindrance to the assay by the
existence of a dialysis section, a LOC with the dialysis section 70 in
the sample input and preparation section can still be used without a loss
of the required functionality.

[0319] Furthermore, the detection section 294 may encompass proteomic
chamber arrays which are identical to the hybridization chamber arrays
but are loaded with probes designed to conjugate or hybridize with sample
target proteins present in non-amplified sample instead of nucleic acid
probes designed to hybridize to target nucleic acid sequences.

[0320] It will be appreciated that the LOC devices fabricated for use in
this diagnostic system are different combinations of functional sections
selected in accordance with the particular LOC application. The vast
majority of functional sections are common to many of the LOC devices and
the design of additional LOC devices for new application is a matter of
compiling an appropriate combination of functional sections from the
extensive selection of functional sections used in the existing LOC
devices.

[0321] Only a small number of the LOC devices are shown in this
description and some more are shown schematically to illustrate the
design flexibility of the LOC devices fabricated for this system. The
person skilled in the art will readily recognise that the LOC devices
shown in this description are not an exhaustive list and many additional
LOC designs are a matter of compiling the appropriate combination of
functional sections.

Sample Types

[0322] LOC variants can accept and analyze the nucleic acid or protein
content of a variety of sample types in liquid form including, but not
limited to, blood and blood products, saliva, cerebrospinal fluid, urine,
semen, amniotic fluid, umbilical cord blood, breast milk, sweat, pleural
effusion, tear, pericardial fluid, peritoneal fluid, environmental water
samples and drink samples. Amplicon obtained from macroscopic nucleic
acid amplification can also be analysed using the LOC device; in this
case, all the reagent reservoirs will be empty or configured not to
release their contents, and the dialysis, lysis, incubation and
amplification sections will be used solely to transport the sample from
the sample inlet 68 to the hybridization chambers 180 for nucleic acid
detection, as described above.

[0323] For some sample types, a pre-processing step is required, for
example semen may need to be liquefied and mucus may need to be
pre-treated with an enzyme to reduce the viscosity prior to input into
the LOC device.

Sample Input

[0324] Referring to FIGS. 1 and 12, the sample is added to the
macroreceptacle 24 of the test module 10. The macroreceptacle 24 is a
truncated cone which feeds into the inlet 68 of the LOC device 301 by
capillary action. Here it flows into the 64 μm wide×60 μm
deep cap channel 94 where it is drawn towards the anticoagulant reservoir
54, also by capillary action.

Reagent Reservoirs

[0325] The small volumes of reagents required by the assay systems using
microfluidic devices, such as LOC device 301, permit the reagent
reservoirs to contain all reagents necessary for the biochemical
processing with each of the reagent reservoirs having a small volume.
This volume is easily less than 1,000,000,000 cubic microns, in the vast
majority of cases less than 300,000,000 cubic microns, typically less
than 70,000,000 cubic microns and in the case of the LOC device 301 shown
in the drawings, less than 20,000,000 cubic microns.

Dialysis Section

[0326] Referring to FIGS. 15 to 21, 33 and 34, the pathogen dialysis
section 70 is designed to concentrate pathogenic target cells from the
sample. As previously described, a plurality of apertures in the form of
3 micron diameter holes 164 in the roof layer 66 filter the target cells
from the bulk of the sample. As the sample flows past the 3 micron
diameter apertures 164, microbial pathogens pass through the holes into a
series of dialysis MST channels 204 and flow back up into the target
channel 74 via 16 μm dialysis uptake holes 168 (see FIGS. 33 and 34).
The remainder of the sample (erythrocytes and so on) stay in the cap
channel 94. Downstream of the pathogen dialysis section 70, the cap
channel 94 becomes the waste channel 72 leading to the waste reservoir
76. For biological samples of the type that generate a substantial amount
of waste, a foam insert or other porous element 49 within the outer
casing 13 of the test module 10 is configured to be in fluid
communication with the waste reservoir 76 (see FIG. 1).

[0327] The pathogen dialysis section 70 functions entirely on capillary
action of the fluid sample. The 3 micron diameter apertures 164 at the
upstream end of the pathogen dialysis section 70 have capillary
initiation features (CIFs) 166 (see FIG. 33) so that the fluid is drawn
down into the dialysis MST channel 204 beneath. The first uptake hole 198
for the target channel 74 also has a CIF 202 (see FIG. 15) to avoid the
flow simply pinning a meniscus across the dialysis uptake holes 168.

[0328] The small constituents dialysis section 682 schematically shown in
FIG. 71 can have a similar structure to the pathogen dialysis section 70.
The small constituents dialysis section separates any small target cells
or molecules from a sample by sizing (and, if necessary, shaping)
apertures suitable for allowing the small target cells or molecules to
pass into the target channel and continue for further analysis. Larger
sized cells or molecules are removed to a waste reservoir 766. Thus, the
LOC device 30 (see FIGS. 1 and 104) is not limited to separating
pathogens that are less than 3 μm in size, but can be used to separate
cells or molecules of any size desired.

Lysis Section

[0329] Referring back to FIGS. 7, 11 and 13, the genetic material in the
sample is released from the cells by a chemical lysis process. As
described above, a lysis reagent from the lysis reservoir 56 mixes with
the sample flow in the target channel 74 downstream of the surface
tension valve 128 for the lysis reservoir 56. However, some diagnostic
assays are better suited to a thermal lysis process, or even a
combination of chemical and thermal lysis of the target cells. The LOC
device 301 accommodates this with the heated microchannels 210 of the
incubation section 114. The sample flow fills the incubation section 114
and stops at the boiling-initiated valve 106. The incubation
microchannels 210 heat the sample to a temperature at which the cellular
membranes are disrupted.

[0330] In some thermal lysis applications, an enzymatic reaction in the
chemical lysis section 130 is not necessary and the thermal lysis
completely replaces the enzymatic reaction in the chemical lysis section
130.

Boiling-Initiated Valve

[0331] As discussed above, the LOC device 301 has three boiling-initiated
valves 126, 106 and 108. The location of these valves is shown in FIG. 6.
FIG. 31 is an enlarged plan view of the boiling-initiated valve 108 in
isolation at the end of the heated microchannels 158 of the amplification
section 112.

[0332] The sample flow 119 is drawn along the heated microchannels 158 by
capillary action until it reaches the boiling-initiated valve 108. The
leading meniscus 120 of the sample flow pins at a meniscus anchor 98 at
the valve inlet 146. The geometry of the meniscus anchor 98 stops the
advancing meniscus to arrest the capillary flow. As shown in FIGS. 31 and
32, the meniscus anchor 98 is an aperture provided by an uptake opening
from the MST channel 90 to the cap channel 94. Surface tension in the
meniscus 120 keeps the valve closed. An annular heater 152 is at the
periphery of the valve inlet 146. The annular heater 152 is
CMOS-controlled via the boiling-initiated valve heater contacts 153.

[0333] To open the valve, the CMOS circuitry 86 sends an electrical pulse
to the valve heater contacts 153. The annular heater 152 resistively
heats until the liquid sample 119 boils. The boiling unpins the meniscus
120 from the valve inlet 146 and initiates wetting of the cap channel 94.
Once wetting the cap channel 94 begins, capillary flow resumes. The fluid
sample 119 fills the cap channel 94 and flows through the valve downtake
150 to the valve outlet 148 where capillary driven flow continues along
the amplification section exit channel 160 into the hybridization and
detection section 52. Liquid sensors 174 are placed before and after the
valve for diagnostics.

[0334] It will be appreciated that once the boiling-initiated valves are
opened, they cannot be re-closed. However, as the LOC device 301 and the
test module 10 are single-use devices, re-closing the valves is
unnecessary.

Incubation Section and Nucleic Acid Amplification Section

[0335] FIGS. 6, 7, 13, 14, 23, 24, 25, 35 to 45, 50 and 51 show the
incubation section 114 and the amplification section 112. The incubation
section 114 has a single, heated incubation microchannel 210 etched in a
serpentine pattern in the MST channel layer 100 from the downtake opening
134 to the boiling-initiated valve 106 (see FIGS. 13 and 14). Control
over the temperature of the incubation section 114 enables enzymatic
reactions to take place with greater efficiency. Similarly, the
amplification section 112 has a heated amplification microchannel 158 in
a serpentine configuration leading from the boiling-initiated valve 106
to the boiling-initiated valve 108 (see FIGS. 6 and 14). These valves
arrest the flow to retain the target cells in the heated incubation or
amplification microchannels 210 or 158 while mixing, incubation and
nucleic acid amplification takes place. The serpentine pattern of the
microchannels also facilitates (to some extent) mixing of the target
cells with reagents.

[0336] In the incubation section 114 and the amplification section 112,
the sample cells and the reagents are heated by the heaters 154
controlled by the CMOS circuitry 86 using pulse width modulation (PWM).
Each meander of the serpentine configuration of the heated incubation
microchannel 210 and amplification microchannel 158 has three separately
operable heaters 154 extending between their respective heater contacts
156 (see FIG. 14) which provides for the two-dimensional control of input
heat flux density. As best shown in FIG. 51, the heaters 154 are
supported on the roof layer 66 and embedded in the lower seal 64. The
heater material is TiAl but many other conductive metals would be
suitable. The elongate heaters 154 are parallel with the longitudinal
extent of each channel section that forms the wide meanders of the
serpentine shape. In the amplification section 112, each of the wide
meanders can operate as separate PCR chambers via individual heater
control.

[0337] The small volumes of amplicon required by the assay systems using
microfluidic devices, such as LOC device 301, permit low amplification
mixture volumes for amplification in amplification section 112. This
volume is easily less than 400 nanoliters, in the vast majority of cases
less than 170 nanoliters, typically less than 70 nanoliters and in the
case of the LOC device 301, between 2 nanoliters and 30 nanoliters.

Increased Rates of Heating and Greater Diffusive Mixing

[0338] The small cross section of each channel section increases the
heating rate of the amplification fluid mix. All the fluid is kept a
relatively short distance from the heater 154. Reducing the channel cross
section (that is the amplification microchannel 158 cross section) to
less than 100,000 square microns achieves appreciably higher heating
rates than that provided by more `macro-scale` equipment. Lithographic
fabrication techniques allow the amplification microchannel 158 to have a
cross sectional area transverse to the flow-path less than 16,000 square
microns which gives substantially higher heating rates. Feature sizes on
the order of 1 micron are readily achievable with lithographic
techniques. If very little amplicon is needed (as is the case in the LOC
device 301), the cross sectional area can be reduced to less than 2,500
square microns. For diagnostic assays with 1,000 to 2,000 probes on the
LOC device, and a requirement of `sample-in, answer out` in less than 1
minute, a cross sectional area transverse to the flow of between 400
square microns and 1 square micron is adequate.

[0339] The heater element in the amplification microchannel 158 heats the
nucleic acid sequences at a rate more than 80 Kelvin (K) per second, in
the vast majority of cases at a rate greater than 100 K per second.
Typically, the heater element heats the nucleic acid sequences at a rate
more than 1,000 K per second and in many cases, the heater element heats
the nucleic acid sequences at a rate more than 10,000 K per second.
Commonly, based on the demands of the assay system, the heater element
heats the nucleic acid sequences at a rate more than 100,000 K per
second, more than 1,000,000 K per second more than 10,000,000 K per
second, more than 20,000,000 K per second, more than 40,000,000 K per
second, more than 80,000,000 K per second and more than 160,000,000 K per
second.

[0340] A small cross-sectional area channel is also beneficial for
diffusive mixing of any reagents with the sample fluid. Before diffusive
mixing is complete, diffusion of one liquid into the other is greatest
near the interface between the two. Concentration decreases with distance
from the interface. Using microchannels with relatively small cross
sections transverse to the flow direction, keeps both fluid flows close
to the interface for more rapid diffusive mixing. Reducing the channel
cross section to less than 100,000 square microns achieves appreciably
higher mixing rates than that provided by more `macro-scale` equipment.
Lithographic fabrication techniques allows microchannels with a cross
sectional area transverse to the flow-path less than 16000 square microns
which gives significantly higher mixing rates. If small volumes are
needed (as is the case in the LOC device 301), the cross sectional area
can be reduced to less than 2500 square microns. For diagnostic assays
with 1000 to 2000 probes on the LOC device, and a requirement of
`sample-in, answer out` in less than 1 minute, a cross sectional area
transverse to the flow of between 400 square microns and 1 square micron
is adequate.

Short Thermal Cycle Times

[0341] Keeping the sample mixture proximate to the heaters, and using very
small fluid volumes allows rapid thermal cycling during the nucleic acid
amplification process. Each thermal cycle (i.e. denaturing, annealing and
primer extension) is completed in less than 30 seconds for target
sequences up to 150 base pairs (bp) long. In the vast majority of
diagnostic assays, the individual thermal cycle times are less than 11
seconds, and a large proportion are less than 4 seconds. LOC devices 30
with some of the most common diagnostic assays have thermal cycles time
between 0.45 seconds to 1.5 seconds for target sequences up to 150 bp
long. Thermal cycling at this rate allows the test module to complete the
nucleic acid amplification process in much less than 10 minutes; often
less than 220 seconds. For most assays, the amplification section
generates sufficient amplicon in less than 80 seconds from the sample
fluid entering the sample inlet. For a great many assays, sufficient
amplicon is generated in 30 seconds.

[0342] Upon completion of a preset number of amplification cycles, the
amplicon is fed into the hybridization and detection section 52 via the
boiling-initiated valve 108.

Hybridization Chambers

[0343] FIGS. 52, 53, 54, 56 and 57 show the hybridization chambers 180 in
the hybridization chamber array 110. The hybridization and detection
section 52 has a 24×45 array 110 of hybridization chambers 180,
each with hybridization-responsive FRET probes 186, heater element 182
and an integrated photodiode 184. The photodiode 184 is incorporated for
detection of fluorescence resulting from the hybridization of a target
nucleic acid sequence or protein with the FRET probes 186. Each
photodiode 184 is independently controlled by the CMOS circuitry 86. Any
material between the FRET probes 186 and the photodiode 184 must be
transparent to the emitted light. Accordingly, the wall section 97
between the probes 186 and the photodiode 184 is also optically
transparent to the emitted light. In the LOC device 301, the wall section
97 is a thin (approximately 0.5 micron) layer of silicon dioxide.

[0344] Incorporation of a photodiode 184 directly beneath each
hybridization chamber 180 allows the volume of probe-target hybrids to be
very small while still generating a detectable fluorescence signal (see
FIG. 54). The small amounts permit small volume hybridization chambers. A
detectable amount of probe-target hybrid requires a quantity of probe,
prior to hybridization, which is easily less than 270 picograms
(corresponding to 900,000 cubic microns), in the vast majority of cases
less than 60 picograms (corresponding to 200,000 cubic microns),
typically less than 12 picograms (corresponding to 40,000 cubic microns)
and in the case of the LOC device 301 shown in the accompanying figures,
less than 2.7 picograms (corresponding to a chamber volume of 9,000 cubic
microns). Of course, reducing the size of the hybridization chambers
allows a higher density of chambers and therefore more probes on the LOC
device. In LOC device 301, the hybridization section has more than 1,000
chambers in an area of 1,500 microns by 1,500 microns (i.e. less than
2,250 square microns per chamber). Smaller volumes also reduce the
reaction times so that hybridization and detection is faster. An
additional advantage of the small amount of probe required in each
chamber is that only very small quantities of probe solution need to be
spotted into each chamber during production of the LOC device.
Embodiments of the LOC device according to the invention can be spotted
using a probe solution volume of 1 picoliter or less.

[0345] After nucleic acid amplification, boiling-initiated valve 108 is
activated and the amplicon flows along the flow-path 176 and into each of
the hybridization chambers 180 (see FIGS. 52 and 56). An end-point liquid
sensor 178 indicates when the hybridization chambers 180 are filled with
amplicon and the heaters 182 can be activated.

[0346] After sufficient hybridization time, the LED 26 (see FIG. 2) is
activated. The opening in each of the hybridization chambers 180 provides
an optical window 136 for exposing the FRET probes 186 to the excitation
radiation (see FIGS. 52, 54 and 56). The LED 26 is illuminated for a
sufficiently long time in order to induce a fluorescence signal from the
probes with high intensity. During excitation, the photodiode 184 is
shorted. After a pre-programmed delay 300 (see FIG. 2), the photodiode
184 is enabled and fluorescence emission is detected in the absence of
the excitation light. The incident light on the active area 185 of the
photodiode 184 (see FIG. 54) is converted into a photocurrent which can
then be measured using CMOS circuitry 86.

[0347] The hybridization chambers 180 are each loaded with probes for
detecting a single target nucleic acid sequence. Each hybridization
chambers 180 can be loaded with probes to detect over 1,000 different
targets if desired. Alternatively, many or all the hybridization chambers
can be loaded with the same probes to detect the same target nucleic acid
repeatedly. Replicating the probes in this way throughout the
hybridization chamber array 110 leads to increased confidence in the
results obtained and the results can be combined by the photodiodes
adjacent those hybridization chambers to provide a single result if
desired. The person skilled in the art will recognise that it is possible
to have from one to over 1,000 different probes on the hybridization
chamber array 110, depending on the assay specification.

Hybridization Chambers with Electrochemiluminescence Detection

[0348] FIGS. 97, 120, 138 and 139 show the hybridization chambers 180 used
in an ECL variant of the LOC device, LOC variant L 729. In this
embodiment of the LOC device, a 24×45 array 110 of hybridization
chambers 180, each with hybridization-responsive ECL probes 237, is
positioned in registration with a corresponding array of photodiodes 184
integrated into the CMOS. In a similar fashion to the LOC devices
configured for fluorescence detection, each photodiode 184 is
incorporated for detection of ECL resulting from the hybridization of a
target nucleic acid sequence or protein with an ECL probe 237. Each
photodiode 184 is independently controlled by the CMOS circuitry 86.
Again, the transparent wall section 97 between the probes 186 and the
photodiode 184 is transparent to the emitted light.

[0349] A photodiode 184 closely adjacent each hybridization chamber 180
allows the amount of probe-target hybrids to be very small while still
generating a detectable ECL signal (see FIG. 97). The small amounts
permit small volume hybridization chambers. A detectable amount of
probe-target hybrid requires a quantity of probe, prior to hybridization,
which is easily less than 270 picograms (corresponding to a chamber
volume of 900,000 cubic microns), in the vast majority of cases less than
60 picograms (corresponding to 200,000 cubic microns), typically less
than 12 picograms (corresponding to 40,000 cubic microns) and in the case
of the LOC device shown in the drawings less than 2.7 picograms
(corresponding to a chamber volume of 9,000 cubic microns). Of course,
reducing the size of the hybridization chambers allows a higher density
of chambers and therefore more probes on the LOC device. In the LOC
device shown, the hybridization section has more than 1,000 chambers in
an area of 1,500 microns by 1,500 microns (i.e. less than 2,250 square
microns per chamber). Smaller volumes also reduce the reaction times so
that hybridization and detection is faster. An additional advantage of
the small amount of probe required in each chamber is that only very
small quantities of probe solution need be spotted into each chamber
during production of the LOC device. In the case of the LOC device shown
in the drawings, the required amount of probe can be spotted using a
solution volume of 1 picoliter or less.

[0350] After nucleic acid amplification, the boiling-initiated valve 108
is activated and the amplicon flows along the flow-path 176 and into each
of the hybridization chambers 180 (see FIGS. 52 and 139). An end-point
liquid sensor 178 indicates when the hybridization chambers 180 are
filled with amplicon so that the heaters 182 can be activated.

[0351] After sufficient hybridization time, the photodiode 184 is enabled
ready for collection of the ECL signal. Then the ECL excitation drivers
39 (see FIG. 105) activate the ECL electrodes 860 and 870 for a
predetermined length of time. The photodiode 184 remains active for a
short time after cessation of the ECL excitation current to maximize the
signal-to-noise ratio. For example, if the photodiode 184 remains active
for five times the decay lifetime of the luminescent emission, then the
signal will have decayed to less than one percent of the initial value.
The incident light on the photodiode 184 is converted into a photocurrent
which can then be measured using CMOS circuitry 86.

Proteomic Assay Chambers

[0352] Some LOC variants, such as LOC variant L 729, are configured to
perform homogeneous protein assays on crude cell lysates within proteomic
assay chamber arrays (see for example 124.1 to 124.3 of FIGS. 116 and
120) for the detection of host cell and/or pathogenic proteins. The
proteomic assay chamber arrays 124.1-124.3 are manufactured and
configured in exactly the same manner as the hybridization chamber arrays
110 (see FIGS. 52, 53, 54 and 56). Each proteomic assay chamber has a
diffusion barrier 175 at the entrance to prevent diffusion of sample and
reagents between chambers, thus avoiding an erroneous result (see FIGS.
84 and 85, which are insets DC and DD of FIG. 81). Where required for
protein hybridization or conjugation, thermal energy is provided by
CMOS-controlled heaters 182 in each chamber. In some embodiments, an
end-point liquid sensor 178 is used to indicate when the proteomic assay
chambers are filled with sample so that the heaters 182 can be activated.
After sufficient time has elapsed, the fluorescent or
electrochemiluminescent signal generated following protein recognition is
detected by the photosensor 44.

Humidifier and Humidity Sensor

[0353] Inset AG of FIG. 6 indicates the position of the humidifier 196.
The humidifier prevents evaporation of the reagents and probes during
operation of the LOC device 301. As best shown in the enlarged view of
FIG. 55, a water reservoir 188 is fluidically connected to three
evaporators 190. The water reservoir 188 is filled with molecular
biology-grade water and sealed during manufacturing. As best shown in
FIGS. 55 and 61, water is drawn into three downtakes 194 and along
respective water supply channels 192 by capillary action to a set of
three uptakes 193 at the evaporators 190. A meniscus pins at each uptake
193 to retain the water. The evaporators have annular shaped heaters 191
which encircle the uptakes 193. The annular heaters 191 are connected to
the CMOS circuitry 86 by the conductive columns 376 to the top metal
layer 195 (see FIG. 37). Upon activation, the annular heaters 191 heat
the water causing evaporation and humidifying the device surrounds.

[0354] The position of the humidity sensor 232 is also shown in FIG. 6.
However, as best shown in the enlarged view of Inset AH in FIG. 58, the
humidity sensor has a capacitive comb structure. A lithographically
etched first electrode 296 and a lithographically etched second electrode
298 face each other such that their teeth are interleaved. The opposed
electrodes form a capacitor with a capacitance that can be monitored by
the CMOS circuitry 86. As the humidity increases, the permittivity of the
air gap between the electrodes increases, so that the capacitance also
increases. The humidity sensor 232 is adjacent the hybridization chamber
array 110 where humidity measurement is most important to slow
evaporation from the solution containing the exposed probes.

Feedback Sensors

[0355] Temperature and liquid sensors are incorporated throughout the LOC
device 301 to provide feedback and diagnostics during device operation.
Referring to FIG. 35, nine temperature sensors 170 are distributed
throughout the amplification section 112. Likewise, the incubation
section 114 also has nine temperature sensors 170. These sensors each use
a 2×2 array of bipolar junction transistors (BJTs) to monitor the
fluid temperature and provide feedback to the CMOS circuitry 86. The CMOS
circuitry 86 uses this to precisely control the thermal cycling during
the nucleic acid amplification process and any heating during thermal
lysis and incubation.

[0356] In the hybridization chambers 180, the CMOS circuitry 86 uses the
hybridization heaters 182 as temperature sensors (see FIG. 56). The
electrical resistance of the hybridization heaters 182 is temperature
dependent and the CMOS circuitry 86 uses this to derive a temperature
reading for each of the hybridization chambers 180.

[0357] The LOC device 301 also has a number of MST channel liquid sensors
174 and cap channel liquid sensors 208. FIG. 35 shows a line of MST
channel liquid sensors 174 at one end of every other meander in the
heated microchannel 158. As best shown in FIG. 37, the MST channel liquid
sensors 174 are a pair of electrodes formed by exposed areas of the top
metal layer 195 in the CMOS structure 86. Liquid closes the circuit
between the electrodes to indicate its presence at the sensor's location.

[0358] FIG. 25 shows an enlarged perspective of cap channel liquid sensors
208. Opposing pairs of TiAl electrodes 218 and 220 are deposited on the
roof layer 66. Between the electrodes 218 and 220 is a gap 222 to hold
the circuit open in the absence of liquid. The presence of liquid closes
the circuit and the CMOS circuitry 86 uses this feedback to monitor the
flow.

Gravitational Independence

[0359] The test modules 10 are orientation independent. They do not need
to be secured to a flat stable surface in order to operate. Capillary
driven fluid flows and a lack of external plumbing into ancillary
equipment allow the modules to be truly portable and simply plugged into
a similarly portable hand held reader such as a mobile telephone. Having
a gravitationally independent operation means the test modules are also
accelerationally independent to all practical extents. They are resistant
to shock and vibration and will operate on moving vehicles or while the
mobile telephone is being carried around.

Dialysis Variants

Leukocyte Target

[0360] The dialysis design described above in the LOC device 301 targets
pathogens. FIG. 59 is a schematic section view of a dialysis section 328
designed to concentrate leukocytes from a blood sample for human DNA
analysis. It will be appreciated that the structure is essentially the
same as that of the pathogen target dialysis section 70 described above
with the exception that apertures in the form of 7.5 micron diameter
holes 165 restrict leukocytes from passing from the cap channel 94 to the
dialysis MST channels 204. In situations where the sample being analysed
is a blood sample, and the presence of haemoglobin from the erythrocytes
interferes with the subsequent reaction steps, addition of an erythrocyte
lysis buffer along with the anticoagulant in the reservoir 54 (see FIG.
22), will ensure that the majority of the lysed erythrocytes (and hence
haemoglobin) will be removed from the sample during this dialysis step. A
commonly used erythrocyte lysis buffer is 0.15M NH4CL, 10 mM
KHCO3, 0.1 mM EDTA, pH 7.2-7.4, but a person skilled in the art will
recognise that any buffer which efficiently lyses erythrocytes can be
used.

[0361] Downstream of the leukocyte dialysis section 328, the cap channel
94 becomes the target channel 74 such that the leukocytes continue as
part of the assay. Furthermore, in this case, the dialysis uptake holes
168 lead to a waste channel 72 so that all smaller cells and components
in the sample are removed. It should be noted that this dialysis variant
only reduces the concentration of the unwanted specimens in the target
channel 74.

[0362] FIG. 72 schematically illustrates a large constituents dialysis
section 686 which also separates any large target constituents from a
sample. The apertures in this dialysis section are fabricated with a size
and shape tailored to withhold the large target constituents of interest
in the target channel for further analysis. As with the leukocyte
dialysis section described above, most (but not all) smaller sized cells,
organisms or molecules flow to a waste reservoir 768. Thus, other
embodiments of the LOC device are not limited to separating leukocytes
that are larger than 7.5 μm in size, but can be used to separate
cells, organisms or molecules of any size desired.

Dialysis Section with Flow Channel to Prevent Trapped Air Bubbles

[0363] Described below is an embodiment of the LOC device referred to as
LOC variant VIII 518 and shown in FIGS. 65, 66, 67 and 68. This LOC
device has a dialysis section that fills with the fluid sample without
leaving air bubbles trapped in the channels. LOC variant VIII 518 also
has an additional layer of material referred to as an interface layer
594. The interface layer 594 is positioned between the cap channel layer
80 and the MST channel layer 100 of the CMOS+MST device 48. The interface
layer 594 allows a more complex fluidic interconnection between the
reagent reservoirs and the MST layer 87 without increasing the size of
the silicon substrate 84.

[0364] Referring to FIG. 66, the bypass channel 600 is designed to
introduce a time delay in the fluid sample flow from the interface waste
channel 604 to the interface target channel 602. This time delay allows
the fluid sample to flow through the dialysis MST channel 204 to the
dialysis uptake 168 where it pins a meniscus. With a capillary initiation
feature (CIF) 202 at the uptake from the bypass channel 600 to the
interface target channel 602, the sample fluid fills the interface target
channel 602 from a point upstream of all the dialysis uptakes 168 from
the dialysis MST channels 204.

[0365] Without the bypass channel 600, the interface target channel 602
still starts filling from the upstream end, but eventually the advancing
meniscus reaches and passes over an uptake belonging to an MST channel
that has not yet filled, leading into air entrapment at that point.
Trapped air reduces the sample flow rate through the leukocyte dialysis
section 328.

Nucleic Acid Amplification Variants

Direct PCR

[0366] Traditionally, PCR requires extensive purification of the target
DNA prior to preparation of the reaction mixture. However, with
appropriate changes to the chemistry and sample concentration, it is
possible to perform nucleic acid amplification with minimal DNA
purification, or direct amplification. When the nucleic acid
amplification process is PCR, this approach is called direct PCR. In LOC
devices where nucleic acid amplification is performed at a controlled,
constant temperature, the approach is direct isothermal amplification.
Direct nucleic acid amplification techniques have considerable advantages
for use in LOC devices, particularly relating to simplification of the
required fluidic design. Adjustments to the amplification chemistry for
direct PCR or direct isothermal amplification include increased buffer
strength, the use of polymerases which have high activity and
processivity, and additives which chelate with potential polymerase
inhibitors. Dilution of inhibitors present in the sample is also
important.

[0367] To take advantage of direct nucleic acid amplification techniques,
the LOC device designs incorporate two additional features. The first
feature is reagent reservoirs (for example reservoir 58 in FIG. 8) which
are appropriately dimensioned to supply a sufficient quantity of
amplification reaction mix, or diluent, so that the final concentrations
of sample components which might interfere with amplification chemistry
are low enough to permit successful nucleic acid amplification. The
desired dilution of non-cellular sample components is in the range of
5× to 20×. Different LOC structures, for example the pathogen
dialysis section 70 in FIG. 4, are used when appropriate to ensure that
the concentration of target nucleic acid sequences is maintained at a
high enough level for amplification and detection. In this embodiment,
further illustrated in FIG. 6, a dialysis section which effectively
concentrates pathogens small enough to be passed into the amplification
section 292 is employed upstream of the sample extraction section 290,
and rejects larger cells to a waste receptacle 76. In another embodiment,
a dialysis section is used to selectively deplete proteins and salts in
blood plasma while retaining cells of interest.

[0368] The second LOC structural feature which supports direct nucleic
acid amplification is design of channel aspect ratios to adjust the
mixing ratio between the sample and the amplification mix components. For
example, to ensure dilution of inhibitors associated with the sample in
the preferred 5×-20× range through a single mixing step, the
length and cross-section of the sample and reagent channels are designed
such that the sample channel, upstream of the location where mixing is
initiated, constitutes a flow impedance 4×-19× higher than
the flow impedance of the channels through which the reagent mixture
flows. Control over flow impedances in microchannels is readily achieved
through control over the design geometry. The flow impedance of a
microchannel increases linearly with the channel length, for a constant
cross-section. Importantly for mixing designs, flow impedance in
microchannels depends more strongly on the smallest cross-sectional
dimension. For example, the flow impedance of a microchannel with
rectangular cross-section is inversely proportional to the cube of the
smallest perpendicular dimension, when the aspect ratio is far from
unity.

Reverse-Transcriptase PCR (RT-PCR)

[0369] Where the sample nucleic acid species being analysed or extracted
is RNA, such as from RNA viruses or messenger RNA, it is first necessary
to reverse transcribe the RNA into complementary DNA (cDNA) prior to PCR
amplification. The reverse transcription reaction can be performed in the
same chamber as the PCR (one-step RT-PCR) or it can be performed as a
separate, initial reaction (two-step RT-PCR). In the LOC variants
described herein, a one-step RT-PCR can be performed simply by adding the
reverse transcriptase to reagent reservoir 62 along with the polymerase
and programming the heaters 154 to cycle firstly for the reverse
transcription step and then progress onto the nucleic acid amplification
step. A two-step RT-PCR could also be easily achieved by utilizing the
reagent reservoir 58 to store and dispense the buffers, primers, dNTPs
and reverse transcriptase and the incubation section 114 for the reverse
transcription step followed by amplification in the normal way in the
amplification section 112.

Isothermal Nucleic Acid Amplification

[0370] For some applications, isothermal nucleic acid amplification is the
preferred method of nucleic acid amplification, thus avoiding the need to
repetitively cycle the reaction components through various temperature
cycles but instead maintaining the amplification section at a constant
temperature, typically around 37° C. to 41° C. A number of
isothermal nucleic acid amplification methods have been described,
including Strand Displacement Amplification (SDA), Transcription Mediated
Amplification (TMA), Nucleic Acid Sequence Based Amplification (NASBA),
Recombinase Polymerase Amplification (RPA), Helicase-Dependent isothermal
DNA Amplification (HDA), Rolling Circle Amplification (RCA), Ramification
Amplification (RAM) and Loop-mediated Isothermal Amplification (LAMP),
and any of these, or other isothermal amplification methods, can be
employed in particular embodiments of the LOC device described herein.

[0371] In order to perform isothermal nucleic acid amplification, the
reagent reservoirs 60 and 62 adjoining the amplification section will be
loaded with the appropriate reagents for the specified isothermal method
instead of PCR amplification mix and polymerase. For example, for SDA,
reagent reservoir 60 contains amplification buffer, primers and dNTPs and
reagent reservoir 62 contains an appropriate nickase enzyme and Exo-DNA
polymerase. For RPA, reagent reservoir 60 contains the amplification
buffer, primers, dNTPs and recombinase proteins, with reagent reservoir
62 containing a strand displacing DNA polymerase such as Bsu. Similarly,
for HDA, reagent reservoir 60 contains amplification buffer, primers and
dNTPs and reagent reservoir 62 contains an appropriate DNA polymerase and
a helicase enzyme to unwind the double stranded DNA strand instead of
using heat. The skilled person will appreciate that the necessary
reagents can be split between the two reagent reservoirs in any manner
appropriate for the nucleic acid amplification process.

[0372] For amplification of viral nucleic acids from RNA viruses such as
HIV or hepatitis C virus, NASBA or TMA is appropriate as it is
unnecessary to first transcribe the RNA to cDNA. In this example, reagent
reservoir 60 is filled with amplification buffer, primers and dNTPs and
reagent reservoir 62 is filled with RNA polymerase, reverse transcriptase
and, optionally, RNase H.

[0373] For some forms of isothermal nucleic acid amplification it may be
necessary to have an initial denaturation cycle to separate the double
stranded DNA template, prior to maintaining the temperature for the
isothermal nucleic acid amplification to proceed. This is readily
achievable in all embodiments of the LOC device described herein, as the
temperature of the mix in the amplification section 112 can be carefully
controlled by the heaters 154 in the amplification microchannels 158 (see
FIG. 14).

[0374] Isothermal nucleic acid amplification is more tolerant of potential
inhibitors in the sample and, as such, is generally suitable for use
where direct nucleic acid amplification from the sample is desired.
Therefore, isothermal nucleic acid amplification is sometimes useful in
LOC variant XLIII 673, LOC variant XLIV 674 and LOC variant XLVII 677,
amongst others, shown in FIGS. 73, 74 and 75, respectively. Direct
isothermal amplification may also be combined with one or more
pre-amplification dialysis steps 70, 686 or 682 as shown in FIGS. 73 and
75 and/or a pre-hybridization dialysis step 682 as indicated in FIG. 74
to help partially concentrate the target cells in the sample before
nucleic acid amplification or remove unwanted cellular debris prior to
the sample entering the hybridization chamber array 110, respectively.
The person skilled in the art will appreciate that any combination of
pre-amplification dialysis and pre-hybridization dialysis can be used.

[0375] Isothermal nucleic acid amplification can also be performed in
parallel amplification sections such as those schematically represented
in FIGS. 64, 69 and 70, multiplexed and some methods of isothermal
nucleic acid amplification, such as LAMP, are compatible with an initial
reverse transcription step to amplify RNA.

Other Design Variants

Flow Rate Sensor

[0376] In addition to temperature and liquid sensors, the LOC device can
also incorporate CMOS-controlled flow rate sensors 740, as schematically
illustrated in FIG. 94 and in LOC Variant X 728 (see FIGS. 76 to 92). The
sensors are used to determine the flow rate in two steps. In the first
step, the temperature of the serpentine heater element 814 is determined
by applying a low current and measuring the voltage to determine the
resistance of the serpentine heater element 814, and therefore the
temperature of the element 814 using the known relationship between
resistance and the temperature of the heater element. At this stage,
minimal heat is being dissipated in the element 814 and the temperature
of the liquid in the channel is equal to the calculated temperature of
the element 814. In the second step, a higher current is applied to the
serpentine heater element 814 such that the temperature of the element
814 increases and some heat is lost to the flowing liquid. By again
measuring the voltage across the element 814 while the higher current is
being applied, the new resistance of the element 814 is determined and
the increased temperature is again calculated by the CMOS circuitry 86.
Using the new temperature of the serpentine heater element 814 and the
known temperature of sample liquid calculated in the first step, the flow
speed of the liquid is determined. From the known channel cross sectional
geometry and the flow speed, the flow rate of the liquid in the channel
is calculated.

Protein Detection Variants

[0377] Some embodiments of the LOC device use a homogeneous protein
detection assay to detect specific proteins within a crude cell lysate.
Numerous homogeneous protein detection assays have been developed for use
in these embodiments of the LOC device. Commonly, these assays utilize
antibodies or aptamers to capture the target protein.

[0378] In one type of assay, an aptamer 141 which binds to a particular
protein 142 is labelled with two different fluorophores or luminophores
143 and 144 which can function as a donor and an acceptor in a
fluorescence resonance energy transfer (FRET) or electrochemiluminescence
resonance energy transfer (ERET) reaction (see FIGS. 108A and 108B). Both
donor 143 and acceptor 144 are linked to the same aptamer 141, and the
change in separation is caused by a change in conformation upon binding
to the target protein 142. For example, an aptamer 141 in the absence of
the target forms a conformation where the donor and acceptor are in close
proximity (see FIG. 108A); upon binding to the target, the new
conformation results in a larger separation between the donor and
acceptor (see FIG. 108B). When the acceptor is a quencher and the donor
is a luminophore, the effect of binding to the target is an increase in
light emission 250 or 862 (see FIG. 108B).

[0379] A second type of assay uses two antibodies 145 or two aptamers 141
that must independently bind to different, non-overlapping epitopes or
regions of the target protein 142 (see FIGS. 109A, 109B, 110A and 110B).
These antibodies 145 or aptamers 141 are labelled with different
fluorophores or luminophores 143 and 144 which can function as a donor
and an acceptor in a fluorescence resonance energy transfer (FRET) or
electrochemiluminescence resonance energy transfer (ERET) reaction. The
fluorophores or luminophores 143 and 144 form part of a pair of short
complementary oligonucleotides 147 attached to the antibodies or aptamers
via long, flexible linkers 149 (see FIGS. 109A and 110A). Once the
antibodies 145 or aptamers 141 bind to the target protein 142, the
complementary oligonucleotides 147 find each other and hybridize to one
another (see FIGS. 109B and 110B). This brings the donors and acceptors
143 and 144 in close proximity to one another resulting in efficient FRET
250 or ERET 862 that is used as a signal for target protein detection.

[0380] To ensure there is no, or very little, background signal as a
result of the oligonucleotides 147 attached to the two antibodies 145 or
aptamers 141 hybridizing to one another in the absence of their binding
to the protein 142, it is necessary to carefully choose the length and
sequence of the complementary oligonucleotides 147 so that the
dissociation constant (kd) for the duplex is relatively high
(˜5 μM). Thus when free antibodies or aptamers labelled with
these oligonucleotides are mixed at nanomolar concentrations, well below
that of their kd, the likelihood of duplex formation and a FRET 250
or ERET 862 signal being generated is negligible. However, when both
antibodies 145 or both aptamers 141 bind to the target protein 142, the
local concentration of the oligonucleotides 147 will be much higher than
their kd resulting in almost complete hybridization and generation
of a detectable FRET 250 or ERET 862 signal.

[0381] The choice of fluorophores and luminophores is an important
consideration when designing a homogeneous protein detection assay. Crude
cell lysates are often turbid and may contain substances which
autofluoresce. In such cases, the use of molecules with long-lasting
fluorescence or electrochemiluminescence and donor-acceptor pairs 143 and
144 which are optimized to give maximal FRET 250 or ERET 862 is desired.
One such pair is europium chelate and Cy5, which has previously been
shown to significantly improve signal-to-background ratio in such a
system when compared with other donor-acceptor pairs, by allowing the
signal to be read after interfering background fluorescence,
electrochemiluminescence or scattered light has decayed. Europium chelate
and AlexaFluor 647 or terbium chelate and Fluorescein FRET or ERET pairs
also work well. The sensitivity and specificity of this approach is
similar to that of enzyme-linked immunosorbent assays (ELISAs), but no
sample manipulation is required.

[0382] In some embodiments of the LOC device, one of the antibodies 145 or
one of the aptamers 141 is attached to the base of the proteomic assay
chamber 124 (see for example FIGS. 116 and 120) and the protein lysate is
combined with the other antibody 145 or aptamer 141 during lysis within
the chemical lysis section 130 to facilitate binding to the first
antibody 145 or aptamer 141 prior to entering the proteomic assay chamber
124. This increases the subsequent speed with which a detectable signal
is generated as only one conjugation or hybridization event is required
within the proteomic assay chamber.

Photodiode

[0383] FIG. 54 shows the photodiode 184 integrated into the CMOS circuitry
86 of the LOC device 301. The photodiode 184 is fabricated as part of the
CMOS circuitry 86 without additional masks or steps. This is one
significant advantage of a CMOS photodiode over a CCD, an alternate
sensing technology which could be integrated on the same chip using
non-standard processing steps, or fabricated on an adjacent chip. On-chip
detection is low cost and reduces the size of the assay system. The
shorter optical path length reduces noise from the surrounding
environment for efficient collection of the fluorescence signal and
eliminates the need for a conventional optical assembly of lenses and
filters.

[0384] Quantum efficiency of the photodiode 184 is the fraction of photons
impinging on its active area 185 that are effectively converted to
photo-electrons. For standard silicon processes, the quantum efficiency
is in the range of 0.3 to 0.5 for visible light, depending on process
parameters such as the amount and absorption properties of the cover
layers.

[0385] The detection threshold of the photodiode 184 determines the
smallest intensity of the fluorescence signal that can be detected. The
detection threshold also determines the size of the photodiode 184 and
hence the number of hybridization chambers 180 in the hybridization and
detection section 52 (see FIG. 52). The size and number of chambers are
technical parameters that are limited by the dimensions of the LOC device
(in the case of the LOC device 301, the dimensions are 1760
μm×5824 μm) and the real estate available after other
functional modules such as the pathogen dialysis section 70 and
amplification section(s) 112 are incorporated.

[0386] For standard silicon processes, the photodiode 184 detects a
minimum of 5 photons. However, to ensure reliable detection, the minimum
can be set to 10 photons. Therefore with the quantum efficiency range
being 0.3 to 0.5 (as discussed above), the fluorescence emission from the
probes should be a minimum of 17 photons but 30 photons would incorporate
a suitable margin of error for reliable detection.

Electrochemiluminescence as an Alternative Detection Method

[0387] Electrochemiluminescence (ECL) involves the generation of species
at electrode surfaces that then undergo electron-transfer reactions to
form excited states that emit light. Electrochemiluminescence differs
from normal chemiluminescence in that formation of the excited species
relies on oxidation or reduction of the luminophore or a coreactant at an
electrode. Coreactants, in this context, are additional reagents added to
the ECL solution which enhance the efficiency of ECL emission. In normal
chemiluminescence, the excited species form purely through mixing of
suitable reagents. The emitting atom or complex is traditionally referred
to as a luminophore. In brief, ECL relies on generating an excited state
of the luminophore, at which point a photon will be emitted. As with any
such process, it is possible for an alternate path to be taken from the
excited state which does not lead to the desired light emission (i.e.
quenching).

[0388] Embodiments of the test module that use ECL instead of fluorescence
detection do not require an excitation LED. Electrodes are fabricated
within the hybridization chambers to provide the electrical pulse for ECL
generation and the photons are detected using the photosensor 44. The
duration and voltage of the electrical pulse are controlled; in some
embodiments, control over the current is used as an alternative to
controlling the voltage.

Luminophore and Quencher

[0389] The ruthenium complex, [Ru(bpy)3]2+, described previously
for use as a fluorescent reporter in the probes, can also be used as a
luminophore in an ECL reaction in the hybridization chambers, with TPrA
(tri-n-propylamine (CH3CH2-CH2)3N) as the coreactant.
Coreactant ECL has the benefit that luminophores are not consumed after
photon emission and the reagents are available for the process to repeat.
Furthermore, the [Ru(bpy)3]2+/TPrA ECL system provides good
signal levels at physiologically relevant conditions of pH in aqueous
solutions. Alternative coreactants which can produce equivalent or better
results than TPrA with ruthenium complexes are N-butyldiethanolamine and
2-(dibutylamino)ethanol.

[0390] FIG. 95 illustrates the reactions occurring during an ECL process
in which [Ru(bpy)3]2+ is the luminophore 864 and TPrA is the
coreactant 866. ECL emission 862 in the [Ru(bpy)3]2+/TPrA ECL
system follows the oxidation of both Ru(bpy)32+ and TPrA at the
anode 860. The reactions are as follows:

Ru(bpy)32+-e.sup.-→Ru(bpy)33+ (1)

TPrA-e.sup.-[TPrA.sup.•].sup.+→TPrA.sup.•+H.sup.+
(2)

Ru(bpy)33++TPrA.sup.•→Ru(bpy)3.sup.•2+-
+products (3)

Ru(bpy)3.sup.•2+→Ru(bpy)32++hv (4)

[0391] The wavelength of the emitted light 862 is around 620 nm and the
anode potential is around 1.1 V with respect to a Ag/AgCl reference
electrode. For the [Ru(bpy)3]2+/TPrA ECL system, either the
Black Hole Quencher, BHQ 2, or Iowa Black RQ described previously, would
be a suitable quencher. In the embodiments described here, the quencher
is a functional moiety which is initially attached to the probe, but
other embodiments are possible in which the quencher is a separate
molecule free in solution.

Hybridization Probes for ECL Detection

[0392] FIGS. 129 and 130 show the hybridization-responsive ECL probes 237.
These are often referred to as molecular beacons and are stem-and-loop
probes, generated from a single strand of nucleic acid, that luminesce
upon hybridization to complementary nucleic acids. FIG. 129 shows a
single ECL probe 237 prior to hybridization with a target nucleic acid
sequence 238. The probe has a loop 240, stem 242, a luminophore 864 at
the 5' end, and a quencher 248 at the 3' end. The loop 240 consists of a
sequence complementary to the target nucleic acid sequence 238.
Complementary sequences on either side of the probe sequence anneal
together to form the stem 242.

[0393] In the absence of a complementary target sequence, the probe
remains closed as shown in FIG. 129. The stem 242 keeps the
luminophore-quencher pair in close proximity to each other, such that
significant resonant energy transfer can occur between them,
substantially eliminating the ability of the luminophore to emit light
after electrochemical excitation.

[0394] FIG. 130 shows the ECL probe 237 in an open or hybridized
configuration. Upon hybridization to a complementary target nucleic acid
sequence 238, the stem-and-loop structure is disrupted, the luminophore
864 and quencher 248 are spatially separated, thus restoring the ability
of the luminophore 864 to emit light. The ECL emission 862 is optically
detected as an indication that the probe has hybridized.

[0395] The probes hybridize with very high specificity with complementary
targets, since the stem helix of the probe is designed to be more stable
than a probe-target helix with a single nucleotide that is not
complementary. Since double-stranded DNA is relatively rigid, it is
sterically impossible for the probe-target helix and the stem helix to
coexist.

Primer-Linked ECL Probes

[0396] Primer-linked stem-and-loop probes and primer-linked linear probes,
otherwise known as scorpion probes, are an alternative to molecular
beacons and can be used for real-time and quantitative nucleic acid
amplification in the LOC device. Real-time amplification is performed
directly in the hybridization chambers of the LOC device. The benefit of
using primer-linked probes is that the probe element is physically linked
to the primer, thus only requiring a single hybridization event to occur
during the nucleic acid amplification rather than separate hybridizations
of the primers and probes being required. This ensures that the reaction
is effectively instantaneous and results in stronger signals, shorter
reaction times and better discrimination than when using separate primers
and probes. The probes (along with polymerase and the amplification mix)
would be deposited into the hybridization chambers 180 during fabrication
and there would be no need for an amplification section on the LOC
device. Alternatively, the amplification section is left unused or used
for other reactions.

Primer-Linked Linear ECL Probes

[0397] FIGS. 131 and 132 show a primer-linked linear ECL probe 693 during
the initial round of nucleic acid amplification and in its hybridized
configuration during subsequent rounds of nucleic acid amplification,
respectively. Referring to FIG. 131, the primer-linked linear ECL probe
693 has a double-stranded stem segment 242. One of the strands
incorporates the primer linked probe sequence 696 which is homologous to
a region on the target nucleic acid 696 and is labelled on its 5' end
with luminophore 864, and linked on its 3' end to an oligonucleotide
primer 700 via an amplification blocker 694. The other strand of the stem
242 is labelled at its 3 end with a quencher molecule 248. After the
initial round of nucleic acid amplification has completed, the probe can
loop around and hybridize to the extended strand with the, now,
complementary sequence 698. During the initial round of nucleic acid
amplification, the oligonucleotide primer 700 anneals to the target DNA
238 (see FIG. 131) and is then extended, forming a DNA strand containing
both the probe sequence and the amplification product. The amplification
blocker 694 prevents the polymerase from reading through and copying the
probe region 696. Upon subsequent denaturation, the extended
oligonucleotide primer 700/template hybrid is dissociated and so is the
double stranded stem 242 of the primer-linked linear probe, thus
releasing the quencher 248. Once the temperature decreases for the
annealing and extension steps, the primer linked probe sequence 696 of
the primer-linked linear ECL probe curls around and hybridizes to the
amplified complementary sequence 698 on the extended strand and light
emission is detected indicating the presence of the target DNA.
Non-extended primer-linked linear ECL probes retain their double-stranded
stem and light emission remains quenched. This detection method is
particularly well suited for fast detection systems as it relies on a
single-molecule process.

Primer-Linked Stem-and-Loop ECL Probes

[0398] FIGS. 133A to 133F show the operation of a primer-linked
stem-and-loop ECL probe 705. Referring to FIG. 133A, the primer-linked
stem-and-loop ECL probe 705 has a stem 242 of complementary
double-stranded DNA and a loop 240 which incorporates the probe sequence.
One of the stem strands 708 is labelled at its 5' end with luminophore
864. The other strand 710 is labelled with a 3'-end quencher 248 and
carries both the amplification blocker 694 and oligonucleotide primer
700. During the initial denaturation phase (see FIG. 133B), the strands
of the target nucleic acid 238 separate, as does the stem 242 of the
primer-linked stem-and-loop ECL probe 705. When the temperature cools for
the annealing phase (see FIG. 133C), the oligonucleotide primer 700 on
the primer-linked stem-and-loop ECL probe 705 hybridizes to the target
nucleic acid sequence 238. During extension (see FIG. 133D), the
complement 706 to the target nucleic acid sequence 238 is synthesized
forming a DNA strand containing both the probe sequence 705 and the
amplified product. The amplification blocker 694 prevents the polymerase
from reading through and copying the probe region 705. When the probe
next anneals, following denaturation (see FIG. 133E), the probe sequence
of the loop segment 240 of the primer-linked stem-and-loop probe (see
FIG. 133F) anneals to the complementary sequence 706 on the extended
strand. This configuration leaves the luminophore 864 relatively remote
from the quencher 248, resulting in a significant increase in light
emission.

ECL Control Probes

[0399] The hybridization chamber array 110 includes some hybridization
chambers 180 with positive and negative ECL control probes used for assay
quality control. FIGS. 134 and 135 schematically illustrate negative
control ECL probes 786 without a luminophore, and FIGS. 136 and 137 are
sketches of positive control ECL probes 787 without a quencher. The
positive and negative control ECL probes have a stem-and-loop structure
like the ECL probes described above. However, an ECL signal 862 (see FIG.
130) will always be emitted from positive control ECL probes 787 and no
ECL signal 862 is ever emitted from negative control ECL probes 786,
regardless of whether the probes hybridize into an open configuration or
remain closed.

[0400] Referring to FIGS. 134 and 135, the negative control ECL probe 786
has no luminophore (and may or may not have a quencher 248). Hence,
whether the target nucleic acid sequence 238 hybridizes with the probe as
shown in FIG. 135, or the probe remains in its stem 242 and loop 240
configuration as shown in FIG. 134, the ECL signal is negligible.
Alternatively, the negative control ECL probe could be designed so that
it always remains quenched. For example, by having an artificial probe
(loop) sequence 240 that will not hybridize to any nucleic acid sequence
within the sample under investigation, the stem 242 of the probe molecule
will re-hybridize to itself and the luminophore and quencher will remain
in close proximity and no appreciable ECL signal will be detected. This
negative control would account for any low level emission that may occur
if the quenching is not complete.

[0401] Conversely, the positive control ECL probe 787 is constructed
without a quencher as illustrated in FIGS. 136 and 137. Nothing quenches
the ECL emission 862 from the luminophore 864 regardless of whether the
positive control probe 787 hybridizes with the target nucleic acid
sequence 238.

[0402] FIGS. 123 and 124 show another possibility for constructing a
positive control chamber. In this case, the calibration chambers 382
which are sealed from the amplicon (or any flow containing target
molecules) can be filled with the ECL luminophore solution such that a
positive signal is always detected at the electrode

[0403] Similarly, the control chambers can be negative control chambers
because the lack of inlets prevents any targets from reaching the probes
such that an ECL signal is never detected.

[0404] FIG. 52 shows a possible distribution of the positive and negative
control probes (378 and 380 respectively) throughout the hybridization
chamber array 110. For ECL, positive and negative control ECL probes 786
and 787 would replace control fluorescent probes 378 and 380,
respectively. The control probes are placed in hybridization chambers 180
along a line extending diagonally across the hybridization chamber array
110. However, the arrangement of the control probes within the array is
arbitrary (as is the configuration of the hybridization chamber array
110).

Calibration Chambers for ECL Detection

[0405] The non-uniformity of the electrical characteristic of the
photodiode 184, response to any ambient light present at the sensor
array, and light originating at other locations in the array, introduce
background noise and offset into the output signal. This background is
removed from each output signal by calibration chambers 382 in the
hybridization chamber array 110 which either do not contain any probes,
contain probes that have no ECL luminophore, or contain probes with a
luminophore and quencher configured such that quenching is always
expected to occur. The number and arrangement of the calibration chambers
382 throughout the hybridization chamber array is arbitrary. However, the
calibration is more accurate if photodiodes 184 are calibrated by a
calibration chamber 382 that is relatively proximate. Referring to FIG.
139, the hybridization chamber array 110 has one calibration chamber 382
for every eight hybridization chambers 180. That is, a calibration
chamber 382 is positioned in the middle of every three by three square of
hybridization chambers 180. In this configuration, the hybridization
chambers 180 are calibrated by a calibration chamber 382 that is
immediately adjacent.

[0406] FIG. 93 shows a differential imager circuit 788 used to substract
the signal from the photodiode 184 corresponding to the calibration
chamber 382 as a result of the applied electrical pulse, from the ECL
signal from the surrounding hybridization chambers 180. The differential
imager circuit 788 samples the signal from the pixel 790 and a "dummy"
pixel 792. Signals arising from ambient light in the region of the
chamber array are also subtracted. The signals from the pixel 790 are
small (i.e. close to dark signal), and without a reference to a dark
level it is hard to differentiate between the background and a very small
signal.

[0407] During use, the "read_row" 794 and "read_row_d" 795 are activated
and M4 797 and MD4 801 transistors are turned on. Switches 807 and 809
are closed such that the outputs from the pixel 790 and "dummy" pixel 792
are stored on pixel capacitor 803 and dummy pixel capacitor 805
respectively. After the pixel signals have been stored, switches 807 and
809 are deactivated. Then the "read_col" switch 811 and dummy "read_col"
switch 813 are closed, and the switched capacitor amplifier 815 at the
output amplifies the differential signal 817.

ECL Levels and Signal Efficiency

[0408] The normal metric of efficiency in ECL is the number of photons
obtained per "Faradaic" electron, i.e. per electron which participates in
the electrochemistry. The ECL efficiency is denoted φECL:

φ ECL = ∫ 0 t I τ ( N A F )
∫ 0 t i τ ( 5 ) ##EQU00001##

where I is the intensity in photons per second, i is the current in
amperes, F is Faraday's constant, and NA is Avogadro's constant.

[0410] The voltage at the working electrode for the
Ru(bpy)32+/TPrA system is approximately +1.1 V (generally
measured in the literature with respect to a reference Ag/AgCl
electrode). Voltages this high shorten electrode lifetimes but this is
not an issue for single-use devices such as the LOC device used in the
present diagnostic system.

[0411] The ideal voltage between the anode and cathode depends on the
combination of solution components and electrode materials. Selecting the
correct voltage can require compromising between the highest signal
levels, reagent and electrode stability, and the activation of undesired
side reactions such as electrolysis of the water in the chamber. In tests
on buffered aqueous Ru(bpy)3]2+/coreactant solution and
platinum electrodes, the ECL emission is maximized at 2.1-2.2 V
(depending on the coreactant choice). Emission intensities drop to
<75% of the peak values for voltages below 1.9 V and above 2.6 V, and
to <50% of the peak values for voltages below 1.7 V and above 2.8 V. A
preferred anode-cathode voltage difference for ECL operation in such
systems is therefore 1.7-2.8 V, with the range 1.9-2.6 V being
particularly preferred. This allows maximization of the emission
intensity as a function of voltage, while avoiding voltages at which
significant gas evolution at the electrodes is observed.

ECL Emission Wavelength

[0412] The wavelength of the emitted light 862 from ECL has an intensity
peak at around 620 nm (measured in air or vacuum), and the emission spans
a relatively broad wavelength range. Significant emission occurs at
wavelengths from around 550 nm to 700 nm. Furthermore, the peak emission
wavelength can vary by ˜10% due to changes in the chemical
environment around the active species. The LOC device embodiments
described here, which incorporate no wavelength-specific filters, have
two advantages for capturing signals with such a broad and variable
spectrum. The first advantage is sensitivity: any wavelength filter
reduces light transmission, even within its pass band, so efficiency is
improved by not including a filter. The second advantage is flexibility:
adjustment of filter pass bands is not required after minor reagent
changes, and the signals are less dependent on minor differences in
non-target components of the input sample.

Solution Volume Participating in ECL

[0413] ECL relies on the availability of luminophore (and coreactant) in
solution. However, as illustrated in FIG. 97, the excited species 868 are
generated only in the solution 872 near the electrodes 860 and 870. The
parameter boundary layer depth in the models presented here, is the depth
of the layer of solution 872 around the electrode 860 in which the
excited species 868 are generated.

[0414] This is a simplification, since solution dynamics can drive the
available concentration upward or downward: [0415] Increased
availability: diffusion and electrophoretic effects will allow exchange
with more of the solution. [0416] Decreased availability: reagents can
adsorb onto the electrodes and may become unavailable to the ECL process.

[0417] For a boundary layer depth value of 0.5 μm, the following
observations are made:

[0418] ECL is observed in experiments where conjugation to magnetic beads
with diameters up to 4.5 μm is used to attract the luminophore 864 to
the anode 860.

[0419] Ru(bpy)32+/TPrA ECL emission 862 as a function of
electrode spacing, for interdigitated electrode arrays, was found to be
maximised at a 0.8 μm electrode spacing. The requirement for a
coreactant 866 in aqueous solutions 872 can be lifted when electrode
spacings are ˜2 μm. This indicates that the excited species 868
diffuse multiple microns, which implies diffusive exchange on a similar
scale for the species in the ground state.

Steady State and Pulsed Operation

[0420] During pulsed activation of the electrodes 860 and 870, the
intensity of the ECL emission 862 (see FIG. 130) is generally higher than
the intensity of the emission 862 from steady-state activation of the
electrodes. Accordingly, the activation signal to the electrodes 860 and
870 is pulse-width modulated (PWM) by the CMOS circuitry 86 (see FIG.
102).

Reagent Recycling and Species Lifetime

[0421] The Ru complex is not consumed in the Ru(bpy)32+/TPrA ECL
system, so the intensity of emission 862 does not reduce with successive
reaction cycles. The lifetime of the rate-limiting step is approximately
0.2 milliseconds giving a total reaction recycling time of approximately
1 millisecond.

Electrophoretic Effects and Other Constraints

[0422] Given the complexity of the solutions in the hybridization chamber,
a large number of phenomena take place when the ECL voltage is turned on.
Electrophoresis of macromolecules, ohmic conduction, and capacitive
effects from small ion migration occur simultaneously.

[0423] Electrophoresis of the oligonucleotides (probes and amplicon) can
complicate the detection of probe-target hybrids, as DNA is highly
negatively charged and attracted to the anode 860. The time scale for
this motion is typically short (in the order of milliseconds).
Electrophoretic effects are strong even though the voltages are moderate
(˜1 V), because the separation between the anode 860 and cathode
870 is small.

[0424] Electrophoresis enhances the ECL emission 862 in some embodiments
of the LOC device and degrades the emission in others. This is addressed
by increasing or decreasing the electrode spacing to get the associated
increases or decreases in electrophoretic effect. Interdigitation of the
anode 860 and the cathode 870 above the photodiode 184 represents the
extreme case of minimizing this separation. Such an arrangement produces
ECL, even in the absence of a coreactant 866 at carbon electrodes 860 and
870.

Ohmic Heating (DC Current)

[0425] The current required to maintain an ECL voltage of ˜2.2 V, is
determined as follows with reference to the ECL cell 874 schematically
illustrated in FIG. 98.

[0426] The DC current through the chamber is determined by two
resistances: the interface resistance Ri between the electrodes 860
and 870 and the bulk of the solution, and the solution resistance Rs
which is derived from the bulk solution resistivity and conduction path
geometry. For solutions with ionic strengths relevant to the conditions
in LOC devices, the chamber resistance is dominated by interfacial
resistances at the electrodes 860 and 870, and Rs can be neglected.

[0427] The effect of the interfacial resistance is estimated by scaling
measurements of macroscopic current flow through similar solutions for
the electrode geometries in the LOC devices.

[0428] Macroscopic measurements of current density through a similar
solution, at platinum electrodes, were taken. Consistent with the
worst-case (high current) approach being taken, overall ionic strength
and ECL reactant concentrations in the test solution were higher than
those used in the LOC devices. The anode area was smaller than the
cathode area, and was surrounded by a cathode with comparable area in a
ring geometry. For an anode consisting of a circle 2 mm in diameter, the
current measured was 1.1 mA, giving a current density of 350 A/m2.

[0429] In the heating model, the electrode area is for the square ring
geometry schematically illustrated in FIG. 98. The anode is a ring with
width 1 μm and thickness 1 μm. The surface area is 196 square
microns, and therefore the calculated current I=69 nA.

[0430] The heating (power=V2/R) was modelled for the worst case in
which all the heat goes into raising the temperature of the water in the
chamber. This leads to heating of chamber contents at 5.8° C./s,
at a voltage difference of 2.2 V, if no allowance for heat removal by the
bulk of the LOC device is made.

[0431] Heating of the chambers by ˜20° C. can cause
denaturation of most hybridization probes. For highly specific probes
intended for mutation detection, it is preferable to further restrict
heating to 4° C. or less. With this level of temperature
stability, single base mismatch-sensitive hybridization, using
appropriately designed sequences, becomes feasible. This allows the
detection of mutations and allelic differences at the level of single
nucleotide polymorphisms. Hence the DC current is applied to the
electrodes 860 and 870 for 0.69 s, to limit the heating to 4° C.

[0432] A current of ˜69 nA passing through the chamber is far more
than can be accommodated as Faradaic current by the ECL species at
micromolar concentrations. Therefore, low-duty-cycle pulsing of the
electrodes 860 and 870 to further reduce heating (to 1° C. or
less) while maintaining sufficient ECL emission 862, does not introduce
complications associated with reagent depletion. In other embodiments,
the current is reduced to 0.1 nA which removes the need for pulsed
activation of the electrodes. Even at currents as low as 0.1 nA, the ECL
emission 862 is luminophore-limited.

Chamber and Electrode Geometry

Maximizing Optical Coupling Between ECL Luminescence and Photosensor

[0433] The immediate chemical precursors of ECL luminescence are generated
within nanometres of the working electrode. Referring again to FIG. 97,
light emission (the excited species 868) generally occurs within microns
or less of that location. Hence the volume immediately adjacent to the
working electrode (anode 860) is visible to the corresponding photodiode
184 of the photosensor 44. Accordingly, the electrodes 860 and 870 are
directly adjacent the active surface area 185 of the corresponding
photodiode 184 in the photosensor 44. Furthermore, the anode 860 is
shaped to increase the length of its lateral periphery `seen` by the
photodiode 184. This aims to maximize the volume of excited species 868
that can be detected by the underlying photodiode 184.

[0434] FIG. 96 schematically illustrates three embodiments of the anode
860. A comb structure anode 878 has the advantage that the parallel
fingers 880 can be interdigitated with the fingers of a cathode 870. The
interdigitated configuration is shown in FIG. 103, and in a partial view
of a LOC layout in FIGS. 120 and 124. The interdigitated configuration
provides a uniform dielectric gap 876 (see FIG. 97) that is relatively
narrow (1 to 2 microns) and the interdigitated comb structure is
relatively simple for the lithographic fabrication process. As discussed
above, a relatively narrow dielectric gap 876 between the electrodes 860
and 870 obviates the need for a coreactant in some solutions 872, as the
excited species 868 will diffuse between anode and cathode. The removal
of the requirement for a coreactant removes the potential chemical impact
of the coreactant on the various assay chemistries and provides a wider
range of possible assay options.

[0435] Referring again to FIG. 96, some embodiments of the anode 860 have
a serpentine configuration 882. To achieve high periphery length while
maintaining tolerance against fabrication errors, it is convenient to
form wide, rectangular meanders 884.

[0436] The anode may have a more complex configuration 886 if necessary or
desirable. For example, it may have a crenulated section 888, a branched
structure 890, or a combination of the two. Partial views of LOC designs
incorporating a branched structure 890 are shown in FIGS. 138 and 139.
The more complicated configurations such as 886 provide a long length of
lateral periphery, and are best suited to solution chemistries where a
coreactant is employed since patterning a closely-spaced opposing cathode
is more difficult.

Electrode Thickness

[0437] Generally, ECL cells involve a planar working electrode which is
viewed externally. Also, traditional microfabrication techniques for
metal layers tend to lead to planar structures with metal thicknesses of
approximately 1 micron. As has been indicated earlier, and shown
schematically in FIGS. 96, 99 and 100, increasing the length of lateral
periphery enhances the coupling between the ECL emission and the
photodiode 184.

[0438] A second strategy to further increase the efficiency of collection
of emitted light 862 (see FIG. 130) by the photodiode 184 is to increase
the thickness of the anode 860. This is shown schematically in FIG. 97.
The part of the participating volume 892 adjacent to the walls of the
working electrode is the region most efficiently coupled to the
photodiode 184. Therefore, for a given width of working electrode 860,
the overall collection efficiency of the emitted light 862 can be
improved by increasing the thickness of the electrodes. Further, since
high current carrying capacity is not required, the width of the working
electrode 860 is reduced as far as is practical. The thickness of the
electrodes 860 and 870 can not increase without restrictions. Noting that
the feature and separation sizes of the electrodes are likely to be of
the order of 1 micron, and that liquid filling makes gaps which are wider
than they are deep unfavourable, the optimum practical thickness for the
electrodes is 0.25 micron to 2 microns.

Electrode Spacing

[0439] The spacing between the electrodes 860 and 870 is important for the
quality of signals in LOC devices, particularly in embodiments where the
electrodes are interdigitated. In embodiments where the anode 860 is a
branched structure such as shown in FIG. 96 and FIG. 100, the spacing
between adjacent elements can also be important. ECL emission efficiency,
and the collection efficiency of the emitted light, should both be
maximised.

[0440] Generation of ECL emission tends to favour electrode spacings on
the order of one micron or less. Small spacings are particularly
attractive when performing ECL in the absence of a coreactant. The fact
that the spacing can be comparable to the wavelength of the emitted light
862 is of limited importance. Therefore, in many embodiments where the
emitted light 862 (see FIG. 130) is measured at a location which does not
require that the light have passed between the electrodes 860 and 870,
making the electrode spacing as small as practical is often the goal. In
embodiments where the emitted light 862 must pass between the electrodes
860 and 870, however, it becomes necessary to move beyond considering
just the ECL emission process, and consider the wave properties of light.

[0441] The wavelength of the emitted light 862 from ECL of
Ru(bpy)32+ is around 620 nm, and therefore 460 nm (0.46
microns) in water. In embodiments where the photodiode 184 and the ECL
excited species 868 are on different sides of the electrode structure,
and the electrode structure is metallic, the emitted light 862 must pass
through a gap between elements of the metallic structures. If this gap is
comparable to the wavelength of the light, diffraction generally reduces
the intensity of propagating light which reaches the photodiode 184. In
cases where the emitted light 862 is incident on the gap at large angles,
however, evanescent mode coupling can be harnessed to improve the
strength of collected signals. Two measures are taken in the LOC devices
to enhance the efficiency of coupling between the photodiode 184 and the
emitted light 862.

[0442] First, the separation between metallic elements is not reduced
below approximately the wavelength of the emitted light in water, i.e.
approximately 0.4 microns. When combined with other observations
regarding small separations between interdigitated electrodes, this
indicates an optimal range for the electrode spacing of 0.4 to 2 microns.

[0443] Second, the distance from the gap between elements to the
photodiode 184 is minimised. In the LOC device embodiments described
here, this indicates that the total thickness of layers between the
electrodes 860 and 870 and the photodiode 184 be one micron or less. In
embodiments where multiple layers are present between the electrodes and
the photodiode, arranging their thicknesses to be quarter-wave or
three-quarter wave layers has the further benefit of suppressing
reflection of the emitted light 862.

Electrode Models

[0444] FIG. 97 is a schematic partial cross-section of the electrodes 860
and 870 in the hybridization chamber. The volume around the lateral
periphery of the anode 860 occupied by the excited species 868, is
sometimes referred to as the participating volume 892. The occluded
region 894 above the anode 860 is ignored because its optical coupling to
the photodiode 184 is negligible.

[0445] A technique for determining whether a particular electrode
configuration provides a foundation for the level of ECL emission 862 for
the underlying photodiode 184 is set out below with reference to FIGS.
98, 99 and 100.

[0446] FIG. 98 is a ring geometry in which the anode 860 is around the
edge of photodiode 184. In FIG. 99, the anode 860 is positioned within
the periphery of the photodiode 184. FIG. 100 shows a more complex
configuration in which the anode 860 has a series of parallel fingers 880
to increase the length of its lateral edges.

[0447] For all of the above configurations, the model calculations are as
follows. For a participating volume 892 of solution VECL, the total
effective number of emitters Nem is:

Nem=Nlumτp/τECL=VECLCLNAτ-
p/τECL (6)

where the participating number of luminophores
Nlum=VECLCLNA, τECL is the lifetime of the
ECL process, CL is the luminophore concentration, τp is the
pulse duration, and NA is Avogadro's number.

[0448] The number of isotropically emitted photons Nphot is:

Nphot=φECLNem (7)

where φECL is the ECL efficiency, defined as the average number
of photons emitted by the ECL reaction of a single luminophore.

[0449] The signal count of electrons, S, from the photodiode is then

S=Nphotφoφq (8)

where φo is the optical coupling efficiency (the number of
photons absorbed by the photodiode 184) and φq is the photodiode
quantum efficiency. The signal is therefore:

S = V ECL C L N A τ p τ ECL φ ECL
φ o φ q ( 9 ) ##EQU00002##

[0450] For FIGS. 98 and 99 electrode configurations, φo is:

φo=(25% photons which are directed towards the photodiode
184)×(10% of photons which are not reflected)

i.e., φo=2.5% for configurations shown in FIGS. 98 and 99

[0451] For the electrode configuration of FIG. 100, 50% of photons are
emitted in a direction pointing towards the photodiode 184, but the
absorption efficiency as a function of angle is unchanged, so

φo=(50% photons which are directed towards the
photodiode)×(10% of photons which are not reflected)

i.e., φo=5% for the configuration of FIG. 100.

[0452] The participating volume 892 depends on the electrode
configuration, and details are presented in the corresponding sections.

[0453] The input parameters for the calculations are listed in the
following:

[0460] This configuration shown in FIG. 101 and FIG. 102 is included as a
limiting case of maximum surface area coupling. In practice, 90% or
better coupling between the electrode surface area and the active surface
area 185 of the photodiode 184 achieves a nearly optimal result, and even
coupling of 50% of the photodiode active surface area 185 to the
electrode surface area provides most of the benefit of the complete
overlay configuration. Complete overlay can be achieved in two
embodiments: first, as indicated schematically in FIG. 101, by employing
a transparent anode 860, in a plane parallel with that of the photodiode
184 and with an area matched to that of the photodiode, and arranging the
anode in immediate proximity to the photodiode 184, such that emitted
light 862 passes through the anode and onto the photodiode. In a second
embodiment shown schematically in FIG. 102, the anode 860 is again
parallel to and registered with the photodiode area, but the solution 872
fills a void between the anode 860 and the photodiode 184. For signal
modelling of a complete overlay configuration, the anode is assumed to be
a complete layer above the photodiode 184, with half of the photons
directed toward the photodiode 184 (absorption efficiency still 10%).

[0461] It is possible to improve the signal and assay beyond the above
models by using surfactants and probe immobilization at the anode.

Maximum Spacing Between ECL Probes and Photodiode

[0462] The on-chip detection of hybridization avoids the needs for
detection via confocal microscopy (see Background of the Invention). This
departure from traditional detection techniques is a significant factor
in the time and cost savings associated with this system. Traditional
detection requires imaging optics which necessarily uses lenses or curved
mirrors. By adopting non-imaging optics, the diagnostic system avoids the
need for a complex and bulky optical train. Positioning the photodiode
very close to the probes has the advantage of extremely high collection
efficiency: when the thickness of the material between the probes and the
photodiode is on the order of 1 micron, the angle of collection of
emission light is up to 174°. This angle is calculated by
considering light emitted from a probe at the centroid of the face of the
hybridization chamber closest to the photodiode, which has a planar
active surface parallel to that chamber face. The cone of emission angles
within which light is able to be absorbed by the photodiode is defined as
having the emitting probe at its vertex and the corner of the sensor on
the perimeter of its planar face. For a 16 micron×16 micron sensor,
the vertex angle of this cone is 170°; in the limiting case where
the photodiode is expanded so that its area matches that of the 28
micron×26.5 micron hybridization chamber, the vertex angle is
174°. A separation between the chamber face and the photodiode
active surface of 1 micron or less is readily achievable.

[0463] Employing a non-imaging optics scheme does require the photodiode
184 to be very close to the hybridization chamber in order to collect
sufficient photons of fluorescence emission. The maximum spacing between
the photodiode and probes is determined as follows.

[0464] Utilizing a ruthenium chelate luminophore and the electrode
configuration of FIG. 100, we calculated 27,000 photons being absorbed by
our 16 micron×16 micron sensor from the respective hybridization
chamber, to generate 8000 electrons assuming a sensor quantum efficiency
of 30%. In performing this calculation we assumed that the
light-collecting region of our hybridization chamber has a base area
which is the same as our sensor area, one quarter of the total number of
the hybridization photons is angled so as to reach the sensor, and a
conservative 10% estimate for the proportion of photons which do not
scatter away from the sensor-dielectric interface. That is, the light
gathering efficiency of the optical system is φ0=0.025.

[0465] More accurately we can write φ0=[(base area of the
light-collecting region of the hybridization chamber)/(photodetector
area)][Ω/4π][10% absorbed], where Ω=solid angle subtended
by the photodetector at a representative point on the base of the
hybridization chamber. For a right square pyramid geometry:

Ω=4 arcsin (a2/(4d02+a2)), where
d0=distance between the chamber and the photodiode, and a is the
photodiode dimension.

[0466] Each hybridization chamber releases 1.1×106 photons. The
selected photodetector has a detection threshold of 17 photons, and for
values of d0 greater than ten times the sensor size (i.e.,
essentially normal incidence) the proportion of photons not reflected at
the sensor surface can be increased from 10% to 90%. Therefore, the
minimum optical efficiency required is:

φ0=17/(1.1×106×0.9)=1.72×10-5

[0467] The base area of the light-emitting region of the hybridization
chamber 180 is 29 micron×19.75 micron.

[0468] Solving for d0, we will get the maximum limiting distance
between the bottom of our hybridization chamber and our photodetector to
be d0=1600 microns. In this limit, the collection cone angle as
defined above is only 0.8°. It should be noted this analysis
ignores the negligible effect of refraction.

LOC Variants

[0469] The LOC device 301 described and illustrated above in full is just
one of many possible LOC device designs. Variations of the LOC device
that use different combinations of the various functional sections
described above will now be described and/or shown as schematic
flow-charts, from sample inlet to detection, to illustrate some of the
combinations possible. The flow-charts have been divided, where
appropriate, into sample input and preparation stage 288, extraction
stage 290, incubation stage 291, amplification stage 292,
pre-hybridization stage 293 and detection stage 294. For all the LOC
variants that are briefly described or shown only in schematic form, the
accompanying full layouts are not shown for reasons of clarity and
succinctness. Also in the interests of clarity, smaller functional units
such as liquid sensors and temperature sensors are not shown but it will
be appreciated that these have been incorporated into the appropriate
locations in each of the following LOC device designs.

LOC Device with ECL Detection

[0470] FIGS. 111 to 127 show a LOC variant 729 with
electrochemiluminescence (ECL) detection. This LOC device prepares 288,
extracts 290, incubates 291, amplifies 292 and detects 294 both human and
pathogen nucleic acids, as well as human and pathogen protein detection.
ECL is used in the hybridization chamber arrays and proteomic assay
chamber arrays for target detection.

[0471] As best shown in FIG. 117, a biological sample (for example, whole
blood) is added to the sample inlet 68. The sample flows through the cap
channel 94 to the anticoagulant surface tension valve 118. The cap 46 is
fabricated with an interface layer 594 positioned between the cap channel
layer 80 and the MST channel layer 100 of the CMOS+MST device 48 (see
FIG. 112). The interface layer 594 allows a more complex fluidic
interconnection between the reagent reservoirs and the MST layer 87
without increasing the size of the silicon substrate 84.

[0472] FIG. 113 shows the MST layer 87 visible on the top surface of the
CMOS+MST device 48. FIG. 114 shows the cap channel layer 80 on the
underside of the cap 46. FIG. 115 superimposes the reservoirs, the cap
channels 94 and the interface channels to illustrate the more
sophisticated plumbing achieved with a cap 46 incorporating an interface
layer 594.

[0473] As best shown in FIG. 117, the interface layer 594 requires the
anticoagulant surface tension valve 118 to have two interface channels
596 and 598. A reservoir-side interface channel 596 connects the
reservoir outlet with the downtakes 92 and a sample-side interface
channel 598 connects the uptakes 96 with the cap channel 94.

[0474] Anticoagulant from the reservoir 54 flows through the MST channels
90 via the reservoir-side interface channel 596 to pin a meniscus at the
uptakes 96. The sample flow along the cap channel 94 dips into the
sample-side interface channel 598 to remove the meniscus so that the
anticoagulant combines with the blood sample as it continues onto the
leukocyte dialysis section 328.

[0475] The leukocyte dialysis section 328 incorporates a bypass channel
600 for filling the flow channel structures without trapped air bubbles
(see FIGS. 117 and 126). The blood sample flows through cap channel 94 to
the upstream end of the large constituents interface channel 730. The
large constituents interface channel 730 is in fluid communication with
the dialysis MST channels 204 via apertures in the form of 7.5 micron
diameter holes 165 (see FIG. 126).

[0476] Referring to FIG. 126, each of the dialysis MST channels 204 lead
from the 7.5 micron diameter holes 165 to respective dialysis uptakes
168. The dialysis uptake holes 168 are open to the small constituents
interface channel 732. However the uptakes are configured to pin a
meniscus rather than allow capillary driven flow to continue. The uptake
belonging to the bypass channel 600 has a capillary initiation feature
202 configured to initiate capillary driven flow into the small
constituents interface channel 732. This ensures the flow begins at the
upstream end of the small constituents interface channel 732 and
sequentially unpins the menisci at the dialysis uptakes 168 as the flow
progresses downstream.

[0477] FIG. 121 shows the downstream end of the leukocyte dialysis section
328. The large constituents interface channel 730 feeds into the large
constituents cap channel 736 and the small constituents interface channel
732 feeds the small constituents cap channel 734. As best shown in FIG.
115, the large constituents cap channel 736 feeds the leukocytes (and any
other large constituents) into the chemical lysis section 130.1 via the
lysis surface tension valve 128.1 where lysis reagent from reservoir 56.1
is added. The chemical lysis section 130.1 has a 3 micron filter downtake
738 at the outlet (see FIG. 117). The filter downtake ensures that no
large constituents reach the lysis chamber exit boiling-initiated valve
206. After sufficient time, the boiling-initiated valve 206 opens the
chemical lysis section 130.1 outlet and the sample flow is split into two
streams. As best shown in FIG. 117, one stream flows to the surface
tension valve 132.1 for the first restriction enzyme, ligase and linker
reservoir 58.1 and the other stream is drawn along a lysed leukocyte
bypass channel 742 directly to the proteomic assay chamber array 124.1 in
the hybridization and detection section 294. Here the sample fills the
proteomic assay chamber array 124.1 (see FIG. 119) containing probes for
hybridization with target human proteins. Probe-target hybrids are
detected with a photosensor 44 (see FIG. 111). The other stream flows
into the leukocyte incubation section 114.1 together with restriction
enzymes, ligase and linker primers from reservoir 58.1.

[0479] The erythrocytes and pathogens from the leukocyte dialysis section
328 are fed to the pathogen dialysis section 70 via the cap channel 734
(see FIGS. 117 and 127). This operates in the same manner as the
leukocyte dialysis section 328 with the exception that the filter
downtakes have 3 micron holes 164 instead of the 7.5 micron holes 165
used for leukocyte dialysis. The erythrocytes remain in the large
constituents interface channel 730 while the pathogens diffuse to the
small constituents interface channel 732.

[0480] FIG. 122 shows the downstream end of the pathogen dialysis section
70. The erythrocytes flow into the large constituents cap channel 736 and
the pathogens fill the small constituents cap channel 734. It will be
appreciated that `large constituents` and `small constituents` are used
in a relative sense as the large constituents output of the pathogen
dialysis section is part of the small constituents output of the
leukocyte dialysis section. The constituents in the large constituents
cap 736 or interface channels are simply larger than the constituents in
the small constituents cap 734 or interface channels within that
particular dialysis section. As best shown in FIGS. 115 and 116, the
erythrocytes in the large constituents cap channel 736 are directed to
the surface tension valve 128.3 for the lysis reagent reservoir 56.3. The
lysis reagent combines with the erythrocytes as the sample fluid fills
the chemical lysis section 130.3. Boiling-initiated valve 206 at the
outlet of the third chemical lysis section 130.3 retains the pathogens
until lysis is complete. When the boiling-initiated valve 206 opens, the
erythrocyte DNA flows directly into the proteomic assay chamber array
124.3 for protein analysis and detection by the photosensor 44 (see FIG.
119).

[0481] The pathogens in the small constituents cap channel 734 are
directed to the surface tension valve 128.2 of the second lysis reagent
reservoir 56.2. The lysis reagent combines with the pathogens as the
sample fluid fills the second chemical lysis section 130.2. After
sufficient time, the boiling-initiated valve 206 opens the chemical lysis
section 130.2 outlet and the sample flow is split into two streams. As
best shown in FIGS. 116 and 118, one stream flows to the surface tension
valve 132.2 for the second restriction enzyme, ligase and linker
reservoir 58.2 and the other stream is drawn along a bypass channel 744
directly to the hybridization and detection section 294. Here the sample
fills the proteomic assay chamber array 124.2 (see FIG. 119) containing
probes for hybridization with target pathogen proteins or other
biomolecules. Probe-target hybrids are detected with the photosensor 44
(see FIG. 111).

[0482] The other stream flows into the pathogen incubation section 114.2
together with restriction enzymes, ligase and linker primers from
reservoir 58.2. After restriction digestion and linker ligation, the
incubator exit valve 207 (also a boiling-initiated valve) opens and flow
continues into the pathogenic DNA amplification section 112.2 (see FIG.
118). As the chamber fills, the amplification mix and polymerase in
reservoirs 60.2 and 62.2 are added via surface tension valves 138.2 and
140.2 respectively. After thermal cycling, the boiling-initiated valve
108 opens for the amplicon to flow into the second hybridization chamber
array 110.2 containing probes for pathogenic DNA targets. Probe-target
hybrids are detected with the photosensor 44 (see FIG. 119).

[0483] Referring to FIG. 120, the hybridization chamber arrays 110.1 and
110.2 and proteomic assay chamber arrays 124.1 to 124.3 have heater
elements 182 made from strips of titanium nitride. There are end-point
liquid sensors 178 that detect when the flow has reached the end of the
hybridization chamber array or proteomic assay chamber array and the
heaters 182 are then activated after a time delay. The flow rate sensor
740 (see FIG. 125) is included in the pathogen incubation section 114.2
to determine the time delay.

[0484] FIGS. 123 and 124 show the calibration chambers 382. They are used
to calibrate the photodiodes 184 to adjust for system noise and
background levels. The photodiode's response and electrical noise
characteristics can vary with location and due to thermal variations. The
output signal from calibration chambers 382, which do not contain any
probes, closely approximates the noise and background in the output
signal from all the chambers. Subtracting the calibration signal from the
output signals generated by the other hybridization chambers
substantially removes the noise and leaves the signal generated by the
electrochemiluminescence (if any). Also, positive and negative control
ECL probes 786 and 787 can be placed in some of the hybridization
chambers 180 for assay quality control.

[0485] Referring to FIG. 116, a humidifier 196, composed of the water
reservoir 188 and evaporators 190, is located in the top left of the
device. The position of the humidity sensor 232 is adjacent to the
hybridization chamber array 110 where humidity measurement is most
important to slow evaporation from the solution containing the exposed
probes.

[0486] By combining the leukocyte and pathogen output dialysis sections,
three output streams are produced (leukocytes, erythrocytes, and
pathogens and other biomolecules) which are processed separately to
enable higher sensitivity and parallel analysis. The output from each
stream is lysed and separately directed to the proteomic assay chamber
arrays for protein detection. The lysed leukocytes and pathogens are also
separately directed to the incubation 114 and amplification 112 sections
for amplification, followed by hybridization for nucleic acid detection.

LOC Device with Thermal Insulation Trench

[0487] As best depicted in FIG. 128, a trench 896 is etched into the back
of the silicon substrate 84. The purpose of the trench is to thermally
insulate the amplification section 112 from the hybridization chamber
array 110. The hybridization array contains detection probes that can
degrade at high temperatures. The trench, when filled with air, has a
thermal conductivity of the order of 6000 times less than that of the
silicon substrate, thereby significantly reducing the heat flux into
adjacent parts of the LOC device.

[0488] This provides two main advantages: an increase in the heating
efficiency in the amplification section 112; and a reduction in the
undesirable temperature rise of the adjacent hybridization section 110.
Improved heating efficiency means less power is required to heat the
amplification section 112 and the temperature reaches its desired
end-point temperature faster and with better spatial uniformity within
the amplification section. A reduction in the temperature rise in the
hybridization section 110 allows for a wider range of probe chemistries
and superior signal quality.

[0489] The trench can be placed around any region on the LOC device to
thermally insulate the components in that region. The width and depth of
the trench 896 are variable to suit the specific application.

CONCLUSION

[0490] The devices, systems and methods described here facilitate
molecular diagnostic tests at low cost with high speed and at the
point-of-care.

The system and its components described above are purely illustrative and
the skilled worker in this field will readily recognize many variations
and modifications which do not depart from the spirit and scope of the
broad inventive concept.